U.S. patent application number 17/506263 was filed with the patent office on 2022-04-21 for integrated chemical/ultrasonic transducer sensor.
The applicant listed for this patent is The Regents of the University of California. Invention is credited to Muyang Lin, Juliane Renata Sempionatto-Moreto, Joseph Wang, Sheng Xu, Lu Yin.
Application Number | 20220117503 17/506263 |
Document ID | / |
Family ID | 1000005956387 |
Filed Date | 2022-04-21 |
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United States Patent
Application |
20220117503 |
Kind Code |
A1 |
Wang; Joseph ; et
al. |
April 21, 2022 |
INTEGRATED CHEMICAL/ULTRASONIC TRANSDUCER SENSOR
Abstract
Disclosed are devices, systems, and methods for multi-modal,
wearable sensors, including an electrochemical-ultrasonic
transducer-based sensor, that can simultaneously detect and monitor
one or more bio-analyte markers and one or more physiological
markers. In some aspects, a wearable, acoustic-electrochemical
sensor device includes a flexible substrate, one or more
electrochemical sensors disposed on the flexible substrate, a
physiological sensor comprising an array of acoustic transducers
disposed on the flexible substrate, wherein the sensor device is
operable to simultaneously detect and monitor one or more analyte
markers and physiological markers including hemodynamic
parameters.
Inventors: |
Wang; Joseph; (La Jolla,
CA) ; Xu; Sheng; (La Jolla, CA) ;
Sempionatto-Moreto; Juliane Renata; (La Jolla, CA) ;
Yin; Lu; (La Jolla, CA) ; Lin; Muyang; (La
Jolla, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The Regents of the University of California |
Oakland |
CA |
US |
|
|
Family ID: |
1000005956387 |
Appl. No.: |
17/506263 |
Filed: |
October 20, 2021 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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63094169 |
Oct 20, 2020 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 5/14546 20130101;
A61B 5/1477 20130101; A61B 5/02055 20130101; A61B 5/02133 20130101;
A61B 5/14532 20130101; A61B 2562/0204 20130101; A61B 5/026
20130101; A61B 5/14521 20130101; A61B 2562/164 20130101; A61B
5/4845 20130101 |
International
Class: |
A61B 5/0205 20060101
A61B005/0205; A61B 5/1477 20060101 A61B005/1477; A61B 5/145
20060101 A61B005/145; A61B 5/021 20060101 A61B005/021; A61B 5/026
20060101 A61B005/026; A61B 5/00 20060101 A61B005/00 |
Claims
1. A wearable, acoustic-electrochemical sensor device, comprising:
a flexible substrate comprising an electrically insulative
material, the flexible substrate capable of attaching and
conforming to skin; an electrochemical sensor comprising two or
more electrodes disposed on the flexible substrate, the two or more
electrodes including a first electrode to detect a signal
associated with an analyte by an electrochemical detection, and a
second electrode configured as a counter electrode or a reference
electrode; a physiological sensor comprising an array of acoustic
transducers disposed on the flexible substrate and a ground wire
coupled to and spanning across each acoustic transducer of the
array, the array of acoustic transducers including an acoustic
transduction material, wherein the physiological sensor is
configured to direct acoustic signals from the array of acoustic
transducers toward a blood vessel in or beneath the skin to detect
a hemodynamic parameter of the blood vessel; and an array of
electrical interconnection structures disposed on the flexible
substrate, wherein at least one of the electrical interconnection
structures is configured as a ground electrical interconnection
structure, and wherein the ground wire of the physiological sensor
spans from the array of acoustic transducers to the ground
electrical interconnection structure, wherein the sensor device is
operable to simultaneously detect and monitor one or more analyte
markers and physiological markers.
2. The sensor device of claim 1, wherein the array of acoustic
transducers of the physiological sensor is spaced apart from the
electrochemical sensor by a distance of at least 0.1 cm.
3. The sensor device of claim 1, wherein the physiological sensor
is configured on a first side of the flexible substrate configured
to attach to the skin, and the electrochemical sensor is configured
on a second side of the flexible sensor opposite to the first side,
such that the electrochemical sensor is able to be exposed to a
biofluid deposited on the electrochemical sensor.
4. The sensor device of claim 1, wherein the physiological sensor
includes a hydrogel material coupled to the array of acoustic
transducers and configured to propagate an acoustic signal
generated at the acoustic transducers to the skin and to propagate
a returned acoustic echo received from the skin to the acoustic
transducers.
5. The sensor device of claim 1, wherein the electrochemical sensor
includes a functionalization layer disposed at least partially on
the first electrode that includes one or more molecules to catalyze
a chemical reaction or bind to the analyte for the electrochemical
detection at the first electrode, and wherein the wearable,
acoustic-electrochemical sensor device further comprises: a second
electrochemical sensor comprising two or more electrodes disposed
on the flexible substrate, the two or more electrodes of the second
electrochemical sensor including a third electrode to detect a
second signal associated with a second analyte by a second
electrochemical detection, and a fourth electrode configured as a
counter electrode or a reference electrode, wherein the second
analyte is different than the analyte detectable at the first
electrode.
6. The sensor device of claim 1, wherein the second electrode is
configured as the reference electrode, and wherein the two or more
electrodes of the electrochemical sensor include a third electrode
configured as the counter electrode.
7. The sensor device of claim 6, wherein the two or more electrodes
of the electrochemical sensor include a fourth electrode configured
as an iontophoresis (IP) electrode, the IP electrode operable to
facilitate extraction of interstitial fluid of the skin or induce
excretion of sweat from the skin.
8. The sensor device of claim 7, wherein the electrochemical sensor
includes a hydrogel coupled to the IP electrode, wherein the
hydrogel entraps one or more chemicals able to cause extraction of
the interstitial fluid or excretion of the sweat upon controlled
release from the hydrogel by an electrical potential applied at the
IP electrode.
9. The sensor device of claim 6, wherein two or more electrodes are
printed electrodes, wherein the first electrode and the counter
electrode comprise a Prussian Blue, and wherein the reference
electrode comprise a silver ink.
10. The sensor device of claim 1, wherein the electrical
interconnection structures are configured as serpentine
interconnection structures that allow for stretching and bending on
the flexible substrate.
11. The sensor device of claim 1, wherein the acoustic transduction
material includes at least one of piezoelectric lead zirconate
titanate (PZT), lead magnesium niobate-lead titanate (PMN-PT), or
polyvinylidene difluoride (PVDF).
12. The sensor device of claim 11, wherein each transducer pixel
includes an aspect ratio of 0.3 or smaller based on a height
dimension to a width dimension, such that aspect ratio is able to
control vibration of the acoustic transduction material to be in a
thickness mode with a particular frequency or frequency range.
13. The sensor device of claim 12, wherein the particular frequency
is 7 MHz; or wherein the frequency range includes 5 MHz to 9
MHz.
14. The sensor device of claim 1, wherein the flexible substrate
includes at least one of a styrene-ethylene-butylene-styrene block
copolymer (SEBS), a styrene-isoprene-styrene block copolymer (SIS),
or a styrene-butylene-styrene block copolymer (SBS).
15. The sensor device of claim 1, wherein the flexible substrate
includes at least one of ECOFLEX.RTM., polydimethylsiloxane (PDMS),
thermoplastic polyurethane (TPU), polyurethane (PU), or
polyethylene vinyl acetate (PEVA).
16. The sensor device of claim 1, wherein the flexible substrate is
structured to include a first substrate layer and a second
substrate layer that is attached to a side of the first substrate
layer, wherein each of the first substrate layer and the second
substrate layer comprises a first region and a second region,
wherein the physiological sensor is coupled to the first region of
the first substrate layer, and the electrochemical sensor is
coupled to the second region of the second substrate layer, wherein
the second substrate layer includes an opening at the first region
such that physiological sensor is exposed through the opening of
the second substrate layer.
17. The sensor device of claim 1, wherein the hemodynamic parameter
includes blood pressure or blood flow.
18. The sensor device of claim 1, further comprising one or more
additional sensors including a temperature sensor, an
electrocardiogram (ECG) sensor, a pressure sensor, or a mechanical
strain sensor.
19. The sensor device of claim 1, wherein the physiological sensor
comprising the acoustic transducers is operable to detect blood
pressure of a user of the wearable, acoustic-electrochemical sensor
device, and wherein the electrochemical sensor is operable to
detect lactate of the user, such that the sensor device is operable
to monitor for septic shock.
20. The sensor device of claim 1, wherein the analyte includes
lactate, cortisol, glucose, alcohol, caffeine, or an electrolyte.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This patent document claims priorities to and benefits of
U.S. Provisional Patent Application No. 63/094,169, titled
"INTEGRATED CHEMICAL/ULTRASONIC TRANSDUCER SENSOR" filed on Oct.
20, 2020. The entire content of the aforementioned patent
application is incorporated by reference as part of the disclosure
of this patent document.
TECHNICAL FIELD
[0002] This patent document relates to sensors including wearable
sensor having electrodes.
BACKGROUND
[0003] Research in bio-sensing has traditionally been restricted to
well-controlled laboratory environments. Such bio-sensing
modalities measure electroencephalogram (EEG), photoplethysmogram
(PPG), pupillometry, eye-gaze and galvanic skin response (GSR) are
typically bulky, require numerous connections, costly, hard to
synchronize, and have low-resolution and poor sampling rates.
Multi-modal bio-sensing has recently shown to be very effective in
affective computing, research in autism, clinical disorders, and
virtual reality among many others. None of the present bio-sensing
systems support multi-modality in a wearable manner outside
controlled laboratory environments with clean, research-grade
measurements. New devices and systems are needed for bio-sensing
applications.
SUMMARY
[0004] Disclosed are devices, systems, and methods for multi-modal,
wearable sensors, including an electrochemical-ultrasonic
transducer-based sensor, that can simultaneously detect and monitor
one or more bio-analyte markers and physiological markers.
[0005] The integration of an electrochemical sensor with ultrasonic
transducer sensor for non-invasive multiplex sensing is disclosed.
The disclosed devices and systems can simultaneously obtain the
epidermal chemical molecular signals and deep-tissue level blood
pressure signals for the detection and monitoring of various
disease symptoms, which cannot be diagnosed based solely on usual
metabolic chemical levels nor abnormal physiological states. The
disclosed integrated sensor is small, light, conformal, wearable
and non-invasive, which also greatly enhance the users' experience
compared to previous sensors.
[0006] The integrated wearable sensor can be applied onto human
epidermis and connected to either instruments or flexible
integrated circuits to intermittently or continuously measure
multiple signals simultaneously. The obtained data can be processed
and translated for users or professionals to interpret. The
chemical sensor section is designed to perform noninvasive
electrochemical, optical, or colorimetric monitoring of biomarkers
present in sweat, interstitial fluid or other epidermal fluid,
including but not limited to glucose, lactate, uric acid and
cortisol. The transducer sensor is designed to measure ultrasonic
echoing signal reflected from arteries and compare with established
correlation with the blood pressure.
[0007] The disclosed technological platform promises large
commercial prospect since it can be easily incorporated as a useful
device for healthcare monitoring in hospitals or at home. The "easy
to wear", "put and forget" and wireless transmission features, make
this wearable sensor a useful device for easy applicability.
Furthermore, this device enables possible real-time monitoring of
interesting target biomarker molecules for clinical diagnostics
combined with blood pressure and heart rate.
[0008] Some example advantages of the disclosed electrochemical and
ultrasound wearable sensors include user-friendliness and fast
diagnostic results.
[0009] In some embodiments in accordance with the present
technology, a wearable, acoustic-electrochemical sensor device
includes a flexible substrate comprising an electrically insulative
material, the flexible substrate capable of attaching and
conforming to skin; an electrochemical sensor comprising two or
more electrodes disposed on the flexible substrate, the two or more
electrodes including a first electrode to detect a signal
associated with an analyte by an electrochemical detection, and a
second electrode configured as a counter electrode or a reference
electrode; a physiological sensor comprising an array of acoustic
transducers disposed on the flexible substrate and a ground wire
coupled to and spanning across each acoustic transducer of the
array, the array of acoustic transducers including an acoustic
transduction material, wherein the physiological sensor is
configured to direct acoustic signals from the array of acoustic
transducers toward a blood vessel in or beneath the skin to detect
a hemodynamic parameter of the blood vessel; and an array of
electrical interconnection structures disposed on the flexible
substrate, wherein at least one of the electrical interconnection
structures is configured as a ground electrical interconnection
structure, and wherein the ground wire of the physiological sensor
spans from the array of acoustic transducers to the ground
electrical interconnection structure, wherein the sensor device is
operable to simultaneously detect and monitor one or more analyte
markers and physiological markers.
BRIEF DESCRIPTION OF THE DRAWINGS
[0010] FIGS. 1A-1C show illustrative diagrams depicting example
embodiments of a wearable integrated acoustic-electrochemical
sensor device, in accordance with the present technology, for
monitoring physiological data, such as blood pressure and heart
rate via ultrasonic transducers, and, in parallel and
non-invasively, monitoring biomarker levels.
[0011] FIG. 1D shows an illustration depicting various different
daily activities inputs and corresponding biomarkers that affect an
individual's body measurable by the disclosed wearable sensor
technology.
[0012] FIG. 1E shows an illustration depicting a layer-by-layer
layout of an example embodiment of a wearable
acoustic-electrochemical sensor device in accordance with the
present technology.
[0013] FIG. 1F shows photos of an example acoustic-electrochemical
sensor device undergoing bending and undergoing stretching.
[0014] FIG. 1G shows a diagram and corresponding data plot
depicting the detection mechanisms of electrochemical sensors for
detection of multiple analytes, for some example embodiments of a
wearable acoustic-electrochemical sensor device in accordance with
the present technology.
[0015] FIG. 1H shows an illustration and data plot depicting an
example signal generation mechanism of an ultrasound transducer of
an example embodiment of a wearable acoustic-electrochemical sensor
device in accordance with the present technology.
[0016] FIG. 1I shows a diagram illustrating an example embodiment
of a flexible biosensor device including physiological sensors and
electrochemical sensors.
[0017] FIGS. 1J and 1K show illustrative diagrams of an example
embodiments of a wearable integrated acoustic-electrochemical
sensor device, in accordance with the present technology.
[0018] FIG. 1L shows a block diagram of an example embodiment of an
electronic device that can electrically interface with an example
wearable acoustic-electrochemical sensor device for various
implementations in accordance with the present technology.
[0019] FIGS. 2A-2Q show diagrams, images and data plots depicting
example implementations characterizing example embodiments of a
multimodal wearable integrated acoustic-electrochemical sensor
device in accordance with the present technology.
[0020] FIGS. 3A-3D show data plots depicting example data for an
on-body evaluation of an example embodiment of a wearable
acoustic-electrochemical sensor device in accordance with the
present technology.
[0021] FIGS. 4A and 4B show data plots depicting example results
from an evaluation of a lactate, glucose, blood pressure sensor
performance and an alcohol, glucose, blood pressure sensor
performance, respectively, using an example embodiment of a
wearable acoustic-electrochemical sensor device in accordance with
the present technology.
[0022] FIGS. 5A and 5B show data plots depicting example results of
an evaluation of continuous lactate, blood pressure, heart rate
sensor performance for an actively fit volunteer and for a
sedentary volunteer, respectively, using an example embodiment of a
wearable acoustic-electrochemical sensor device in accordance with
the present technology.
[0023] FIG. 6 shows a diagram illustrating a fabrication method for
layer-by-layer printing and assembling of an integrated sensor, in
accordance with the present technology.
[0024] FIG. 7 shows a diagram illustrating a method for assembly
ultrasound transducers for example embodiments of a wearable
integrated acoustic-electrochemical sensor device, in accordance
with the present technology.
[0025] FIG. 8 shows a diagram and an image depicting example
transducer dimensions and conformability for example embodiments of
a wearable integrated acoustic-electrochemical sensor device, in
accordance with the present technology.
[0026] FIG. 9 shows images depicting adhesion of example
piezoelectric lead zirconate titanate (PZT) transducers to a
substrate of an example embodiment of a wearable integrated
acoustic-electrochemical sensor device.
[0027] FIG. 10 is a diagram illustrating electrochemical sensor
modifications and reaction mechanisms for example embodiments of a
wearable integrated acoustic-electrochemical sensor device, in
accordance with the present technology.
[0028] FIG. 11 shows data plots depicting an in vitro
characterization of a lactate sensor in example implementations of
a wearable integrated acoustic-electrochemical sensor device.
[0029] FIG. 12 shows data plots depicting an in vitro
characterization of a glucose sensor in example implementations of
a wearable integrated acoustic-electrochemical sensor device.
[0030] FIG. 13 shows data plots depicting an in vitro
characterization of an alcohol sensor in example implementations of
a wearable integrated acoustic-electrochemical sensor device.
[0031] FIG. 14 shows data plots depicting an in vitro
characterization of a caffeine sensor in example implementations of
a wearable integrated acoustic-electrochemical sensor device.
[0032] FIG. 15 shows data plots depicting ultrasound transducer
characterization on phantom in example implementations of a
wearable integrated acoustic-electrochemical sensor device.
[0033] FIG. 16 is a diagram and a data plot illustrating optimal
channel selection for accurate artery diameter tracking in example
implementations of a wearable integrated acoustic-electrochemical
sensor device.
[0034] FIG. 17 shows data plots depicting electrode electrochemical
stability under repeated stretching tests in example
implementations of a wearable integrated acoustic-electrochemical
sensor device.
[0035] FIG. 18 show data plots depicting sensor electrochemical
stability under repeated stretching tests in example
implementations of a wearable integrated acoustic-electrochemical
sensor device.
[0036] FIG. 19 shows images depicting structural integrity of a
stretchable silver and PB/carbon ink composites in example
implementations of a wearable integrated acoustic-electrochemical
sensor device.
[0037] FIG. 20 shows images and data plots illustrating
electrochemical performance under mechanical deformation in example
implementations of a wearable integrated acoustic-electrochemical
sensor device.
[0038] FIG. 21 shows images and data plots illustrating the
electrochemical performance of an example GOx modified biosensor
under mechanical deformation in example implementations of a
wearable integrated acoustic-electrochemical sensor device.
[0039] FIG. 22 shows a data plot depicting the BP signal measured
on-body while turning the neck 90.degree. to the side, with no
obvious change of signal quality, conducted in example
implementations of an example embodiment of a wearable integrated
acoustic-electrochemical sensor device.
[0040] FIG. 23 shows data plots depicting the BP variation during
the Valsalva maneuver, conducted in example implementations of an
example embodiment of a wearable integrated
acoustic-electrochemical sensor device.
[0041] FIG. 24 shows data plots depicting standard additions to
determine caffeine concentration in sweat in example
implementations.
[0042] FIG. 25 shows data plots depicting a reversibility test for
an example voltammetric caffeine sensor in example implementations
of a wearable integrated acoustic-electrochemical sensor
device.
[0043] FIG. 26 shows diagrams illustrating an example embodiment of
a fabrication method, in accordance with the present technology,
for preparing and assembling of hydrogel layers.
[0044] FIG. 27 shows photos depicting an example implementation of
assembly of example iontophoretic sensors and ultrasound
hydrogels.
[0045] FIG. 28 shows photos depicting an example implementation of
a transfer process of an example wearable acoustic-electrochemical
sensor device.
[0046] FIG. 29 shows a diagram and data plots depicting
characterization of an example multimodal wearable sensor in
example implementations.
[0047] FIG. 30 depicts diagrams and data plots illustrating on body
cross-talking evaluation of the example multimodal wearable sensor
in example implementations.
[0048] FIG. 31 shows diagrams and data plots depicting an in vitro
cross-talking evaluation of the example multimodal wearable sensor
in example implementations.
[0049] FIG. 32 shows images and data plots depicting preparation
and test results of an example solid ultrasound gel in example
implementations.
[0050] FIG. 33 shows a diagram and data plots depicting the effect
of an example embodiment of a substrate on ultrasound
transmission.
[0051] FIG. 34 shows data plots depicting an on-body evaluation of
an example wearable acoustic-electrochemical sensor device for
caffeine intake.
[0052] FIG. 35 shows diagrams and data plots illustrating example
implementations of on-body caffeine detection and pH variation.
[0053] FIG. 36 shows data plots illustrating example control
experiments of electrochemical sensing without a sensing
recognition layer.
[0054] FIG. 37 shows data plots depicting example control
experiments for characterizing response for lactate and glucose
recordings without exercise and food ingestion.
[0055] FIG. 38 shows data plots depicting example control
experiments for characterizing response for alcohol and glucose
recordings without alcohol and food ingestion.
[0056] FIG. 39A shows a diagram depicting the design of an example
embodiment of a wearable, integrated acoustic-electrochemical
sensor device, in accordance with the present technology, for the
simultaneous monitoring of blood pressure along with sweat alcohol,
caffeine and lactate, and ISF glucose chemical markers.
[0057] FIGS. 39B-39G shows diagrams and data plots depicting the
design of an example embodiment of a wearable, integrated
acoustic-electrochemical sensor device, in accordance with the
present technology, for the simultaneous monitoring of blood
pressure along with sodium and lactate from sweat and glucose from
ISF for continuous, simultaneous monitoring of sepsis.
[0058] FIG. 40 shows an image and a diagram depicting a design of
an example embodiment of a stretchable integrated blood
pressure-electrochemical sensing patch device in accordance with
the present technology for the simultaneous detection of sweat
sodium and lactate, and blood pressure.
[0059] FIG. 41 shows data plots depicting an in vitro
characterization of a sodium sensor in example implementations of a
wearable integrated acoustic-electrochemical sensor device.
[0060] FIG. 42 shows data plots depicting an in vitro
characterization of a sodium sensor in example implementations of a
wearable integrated acoustic-electrochemical sensor device.
[0061] FIG. 43 shows a data plot depicting an in vitro
characterization of a lactate sensor in the presence of sodium in
example implementations of a wearable integrated
acoustic-electrochemical sensor device.
[0062] FIG. 44 shows data plots depicting an example implementation
of continuous sodium/lactate//BP/HR performance.
[0063] FIG. 45 shows data plots depicting an example implementation
of continuous alcohol monitoring in stimulated sweat for two
volunteers.
[0064] FIG. 46 shows data plots depicting correlation curves for
sweat and ISF analytes in example implementations.
DETAILED DESCRIPTION
[0065] Intertwined with concepts of telehealth, the internet of
medical things, and precision medicine, wearable sensors offer
attractive features to actively and remotely monitor physiological
parameters. Wearable sensors can generate data continuously without
causing any discomfort or interruptions to daily activity, thus
enhancing wearer's self-monitoring compliance and improving patient
care quality. Wearable sensors can be used for the monitoring of
single physical parameters, such as the electrocardiogram (ECG) and
blood pressure (BP), and of biochemical parameters, such as
glucose.
[0066] Recent efforts have led to the integration of physical and
chemical sensors into a single wearable device, such as sensors for
ECG with lactate or glucose for monitoring athlete's performance,
and temperature with metabolites and electrolytes for signal
calibration. Yet, to the best of our knowledge, an in-depth study
of the correlation of cardiovascular parameters, particularly blood
pressure, with biomarker levels using an integrated hybrid wearable
sensor remains unexplored.
[0067] Blood pressure and heart rate (HR), two of the most
important vital signs, can dynamically and directly reflect the
physiological status of the body. These cardiovascular parameters
can be affected by fluctuations of various biomarker concentrations
originated from activities, such as movement, stress, or intake of
food, drinks, and drugs, that can lead to sudden, sometimes lethal
alterations. Multimodal BP-chemical sensing could thus have
tremendous clinical value, especially for people with underlying
health conditions, such as the elderlies, obese individuals,
diabetic and cardiovascular patients, as their physiological
response to normal day-to-day activities might differ from healthy
people. Further, the prevention, diagnosis, and treatment of many
diseases can greatly benefit from the simultaneous monitoring of
cardiovascular parameters and biomarker levels. These include acute
and deadly septic shock, which commonly involves sudden drops in BP
accompanied by rapidly increasing blood lactate levels and
hypo/hyperglycemia-induced hypo/hypertension which increases the
risks of stroke, cardiac diseases, retinopathy, and nephropathy in
diabetic patients. Simultaneous tracking of metabolites and
hemodynamic parameters using the same device can increase the
patient compliance towards self-monitoring, as it obviates the need
of using multiple devices for measuring these parameters, towards
preventing dangerous cardiac events and saving valuable lives. The
combination of transdisciplinary sensing modalities into a single
miniaturized skin conformal wearable platform can yield significant
additional advantages.
[0068] As an example, tiny critically ill and premature infants
need continuous monitoring of various dangerous conditions, ranging
from hypoglycemia and sepsis-like infection to open-heart surgeries
where blood pressure and lactate or glucose need to be monitored
continuously. Current neonate monitoring platforms require the
application of multiple, often invasive, sensors coupled to bulky
instruments on their tiny body that pose severe injury risks and
barriers to parent-baby bonding. By integrating different sensing
modalities on a single flexible, skin-worn tattoo-like patch,
vulnerable patients--from neonates to elderly--can leverage their
monitoring device with minimal discomfort or obtrusiveness.
Furthermore, the recent global pandemic has also highlighted the
urgent needs for remote self-monitoring devices, with particular
attention to the management of high BP and diabetes, which are
major factors in the deaths of COVID-19 patients. A comprehensive
cardiovascular/biomarker self-monitoring platform would enhance
users' self-awareness to their health conditions, and alert them
and their caregivers to the occurrence of abnormal physiological
changes.
[0069] Disclosed are devices, systems, and methods for multi-modal,
wearable sensors, including an electrochemical-ultrasonic
transducer-based sensor, that can simultaneously detect and monitor
one or more bio-analyte markers and one or more physiological
markers.
[0070] In some example embodiments, a conformal, stretchable, and
integrated wearable sensor is disclosed that can simultaneously
monitor blood pressure, heart rate, and levels of glucose, lactate,
caffeine, and alcohol, toward dynamic and comprehensive health
self-monitoring. The example conformal, stretchable, and integrated
wearable sensor can use ultrasonic transducers for monitoring the
BP and HR, and electrochemical sensors for measuring the levels of
biomarkers.
[0071] The growing demand for tracking the effects of diverse daily
activities upon the body's physiological response calls for
simultaneous tracking of metabolic and hemodynamic parameters on a
single wearable device. Implementations of the disclosed technology
present the first demonstrations of an integrated wearable sensor
that monitors the blood pressure and heart rate via ultrasonic
transducers, along with parallel non-invasive electrochemical
detection of biomarker levels, such as glucose, lactate, caffeine,
and alcohol, in sweat and interstitial fluid. Such simultaneous
non-invasive blood-pressure/chemical sensing was implemented by
monitoring the dynamic effects of everyday activities, such as
exercise and intake of food and drinks, upon the user's
physiological states. For example, by leveraging novel material
selection and assembly processes, the multiplexed sensing
modalities were optimized in some implementations to ensure
reliable sensing without crosstalk between individual sensors,
along with mechanical resiliency and flexibility for conformal
contact to curved skin surfaces. The simultaneous acoustic and
electrochemical sensors were evaluated on multiple human subjects
under different stimuli, and the dynamic correlation of the
hemodynamic activities and corresponding metabolic variations was
monitored and discussed. Such multimodal blood-pressure/chemical
wearable sensor offers a collection of previously unavailable
information towards enhancing our understanding of the body's
response to common activities, while holding considerable promise
for predicting abnormal cardiac events and improving remote,
telemetric, and personalized healthcare medical outcome.
[0072] The technology disclosed in this patent document can be
implemented for various sensing applications, including, for
example, methods, devices and systems for integrated, non-invasive,
wearable multiplex epidermal sensing. In some embodiments, the
disclosed sensor system contains two or more components, including
at least one electrochemical sensor for the sensing of sodium,
potassium, glucose, lactate, cortisol or other chemicals in human
sweat or interstitial fluids; and at least one ultrasound
transducer sensor for the sensing of human physiological signals
such as blood pressure, heart rate, and other physiological
signals. Other types of sensors include temperature sensor, ECG
sensors, pressure sensors or strain sensors can also be integrated.
Such integrated electrochemical-ultrasound sensors offer
comprehensive diagnosis of various symptoms based on both the
chemical signals and the physiological signals of the patient in a
non-invasive manner.
[0073] In some example embodiments disclosed herein, a wearable
sensor device includes an ultrasound device and an electrochemical
sensor device for detecting a variety of physiological parameters
associated with a response or condition of a user wearing the
device, e.g., including monitoring septic shock. Such example
embodiments of the disclosed wearable sensors may be referred to
herein as an acoustic-chem sensor. The wearable acoustic-chem
sensor may include: (i) a blood pressure (BP) sensor provided by
the ultrasound device contingent, and (ii) an electrochemical
sensor. Various example embodiments and implementations of a
wearable acoustic-chem sensor in accordance with the present
technology are described further below.
[0074] In some aspects, the growing demand for tracking the effects
of diverse daily activities upon the body's physiological response
calls for simultaneous tracking of metabolic and hemodynamic
parameters on a single wearable device. In some embodiments in
accordance with the disclosed technology, a wearable
ultrasonic-electrochemical integrated sensor is configured to
monitor the blood pressure and heart rate via ultrasonic
transducers, along with parallel non-invasive electrochemical
detection of biomarker levels, e.g., including but not limited to
glucose, lactate, caffeine, and alcohol, in sweat and interstitial
fluid. Such simultaneous non-invasive blood-pressure/chemical
sensing was employed in example implementations by monitoring the
dynamic effects of everyday activities, such as exercise and intake
of food and drinks, upon the user's physiological states.
Leveraging novel material selection and assembly processes, the
multiplexed sensing modalities were optimized to ensure reliable
sensing without crosstalk between individual sensors, along with
mechanical resiliency and flexibility for conformal contact to
curved skin surfaces. The simultaneous acoustic and electrochemical
sensors were evaluated on multiple human subjects under different
stimuli, and the dynamic correlation of the hemodynamic activities
and corresponding metabolic variations was monitored and discussed.
Such multimodal blood-pressure/chemical wearable sensor offers a
collection of previously unavailable information towards enhancing
our understanding of the body's response to common activities,
while holding considerable promise for predicting abnormal cardiac
events and improving remote, telemetric, and personalized
healthcare medical outcome.
[0075] FIGS. 1A-1C show illustrative diagrams depicting an example
embodiment of a wearable integrated acoustic-electrochemical sensor
device 100 configured to attach to skin of a user and
simultaneously monitor at least one physiological parameter and
electrochemical parameter of the user.
[0076] FIG. 1A shows a block diagram of an example embodiment of
the wearable integrated acoustic-electrochemical sensor device 100.
The wearable acoustic-chem sensor device 100 includes a flexible
substrate 101 comprising an electrically insulative material, in
which the flexible substrate 101 is bendable and/or stretchable and
capable of attaching and conforming to skin. The wearable
acoustic-chem sensor device 100 includes one or more physiological
sensors 110 and one more electrochemical sensors 120, which are
configured on and/or in the flexible (bendable and/or stretchable)
substrate 101.
[0077] In various embodiments of the wearable acoustic-chem sensor
device 100, the one or more electrochemical sensors 120 can include
two or more electrodes disposed on the flexible substrate 101, in
which the two or more electrodes include a first electrode 126A to
detect a first signal associated with a first analyte by an
electrochemical detection, and a second electrode 126B configured
as a counter electrode or a reference electrode to detect a second
signal. In some embodiments, for example, the two or more
electrodes of the one or more electrochemical sensors 120 can
include optionally an iontophoretic electrode 129 to facilitate
extraction of a biofluid, e.g., such as ISF, and/or induce
excretion of a biofluid, e.g., such as sweat. In various
implementations, for example, the one or more electrochemical
sensors 120 can be configured to detect the first signal through
sensing the first analyte by an electrochemical detection,
including, but not limited to: an enzymatic reaction for
electrochemical detection; a non-enzymatic catalytic reaction
(e.g., involving a non-biological catalyst material) for
electrochemical detection; a molecularly-imprinted polymer
facilitated reaction for electrochemical detection; an aptamer
reaction for electrochemical detection; an antibody reaction for
electrochemical detection; an ion-selective membrane facilitated
reaction for electrochemical detection; and/or potential-specific
redox reaction. In some implementations, for example, the one or
more electrochemical sensors 120 can be configured to sensing the
analyte through specific or non-specific adsorption of the analyte
at the detecting electrode for analyte detection. In some
embodiments, for example, the electrodes of the one or more
electrochemical sensors 120 are configured as surface electrodes,
which can include various shapes and sizes, including circular,
elliptical, square, rectangular, triangular, or other shapes. In
some embodiments, for example, the one or more electrochemical
sensors 120 can be configured to detect multiple analytes
simultaneously from one or more biofluids in contact with
electrodes of the one or more electrochemical sensors, where the
analytes include but are not limited to: glucose, lactate, cortisol
and/or other hormones, alcohol, caffeine, urea, uric acid,
acetaminophen, a pharmaceutically-prescribed drug or substance
(e.g., blood pressure regulating drug, L-DOPA, etc.), an illicit
drug or substance (e.g., an opioid, THC, etc.), a vitamin, or an
electrolyte including sodium, chloride, fluoride, magnesium, zinc,
or potassium. The biofluids can include, but are not limited to,
interstitial fluid, sweat, blood, urine, tears, etc.
[0078] In various embodiments of the wearable acoustic-chem sensor
device 100, the one or more physiological sensors 110 can include
an ultrasound sensor that comprises an array of acoustic
transducers 116 disposed on the flexible substrate 101 and a ground
wire 119 coupled to and spanning across each acoustic transducer of
the array, of which the array of acoustic transducers include an
acoustic transduction material. In various implementations, for
example, the one or more physiological sensors 110 is configured to
direct acoustic signals from the array of acoustic transducers
toward a blood vessel in or beneath the skin to detect a
hemodynamic parameter of the blood vessel. In some embodiments, for
example, the one or more physiological sensors 110 can optionally
include an electrocardiogram sensor comprising at least two
electrodes disposed on the flexible substrate 101 to measure a
bioelectrical potential across an area of the user's skin over the
user's heart.
[0079] The wearable acoustic-chem sensor device 100 can include an
array of electrical interconnection 105 structures disposed on the
flexible substrate 101, in which at least one of the electrical
interconnection structures 105 is configured as a ground electrical
interconnection structure that connects to the ground wire 119 of
the one or more physiological sensors 110, e.g., the ground wire
119 spans from the array of acoustic transducers 116 to the ground
electrical interconnection structure. In some embodiments, for
example, the wearable acoustic-chem sensor device 100 can include
an array of electrical contact sites 107 that are coupled to the
electrical interconnection structures 105. In some embodiments, for
example, the wearable acoustic-chem sensor device 100 can be
interfaced to an analytical device comprising a circuit and/or
processing unit, e.g., by the electrical contact sites 107, e.g.,
in which a contact site can be electrically coupled to an electrode
or other electrical component of the sensor(s) 110 and sensor(s)
120 via the electrical interconnection wires 105. The diagrams of
FIG. 1B and FIG. 1C show an example design and mechanism of an
example embodiment of the wearable acoustic-chem sensor device 100
configured as a stretchable integrated blood pressure-chemical
sensing patch device.
[0080] FIG. 1B shows an illustration of an example embodiment of
the wearable acoustic-chem sensor device 100, shown as wearable
acoustic-chem sensor device 100B, attached to the skin of the body,
e.g., on a user's neck. The example wearable acoustic-chem sensor
device 100B is configured to simultaneously monitor the user's
blood pressure and heart rate via the one or more physiological
sensors 110B, e.g., comprising ultrasonic transducers, configured
in parallel with the one or more electrochemical sensors 120B to
non-invasively and electrochemically detect biomarker levels from
biofluids, such as sweat and interstitial fluid (ISF). The
illustration of FIG. 1B shows an electrochemical detection of a
first analyte in ISF measurable at a first electrode contingent
127B of an example electrochemical sensor 120B and an
electrochemical detection of a second analyte in sweat measurable
at a second electrode contingent 128B of another example
electrochemical sensor 120B of the example wearable acoustic-chem
sensor device 100B shown in FIG. 1B. Each of the first and second
sensor contingents 127B and 128B include four individual electrodes
structured to detect a target analyte in the ISF and sweat,
respectively. For example, each of the first and second sensor
contingents 127B and 128B include a detecting electrode that can be
modified with a functionalization layer to facilitate a reaction
involving the target analyte to detect a parameter of the target
analyte in the ISF and sweat respectively (which is shown in FIG.
1B as a circular electrode, and with a caption illustrating an
electrochemical reactions); and each of the first and second sensor
contingents 127B and 128B include a reference electrode, a counter
electrode, and an iontophoretic electrode. In some embodiments, for
example, the detecting electrodes can be configured with a circular
geometry and having a diameter of 1 cm, like that shown in the
particular example illustrated in FIG. 1B; yet, it is understood
that the electrodes of the wearable acoustic-chem sensor device 100
can be configured in a variety of shapes and sizes to suit the
desired purpose and application. In some embodiments, for example,
the electrical interconnection structures 105B can be configured as
serpentine interconnection wires, which allow for stretching and
bending on the flexible substrate 101B.
[0081] FIG. 1C shows an illustration depicting two diagrams of the
example embodiment of a wearable acoustic-chem sensor device 100B.
In the left diagram, the example wearable acoustic-chem sensor
device 100B's acoustic transducers 110B configured for blood
pressure sensing and electrochemical sensors 120B include a
hydrogel for sweat stimulation and ISF extraction, respectively. In
the right diagram, the wearable acoustic-chem sensor device 100B is
shown attached to skin for acoustic sensing and implementing
iontophoresis mechanism of the integrated sensor. For example, the
acoustic transducers 110B of the sensor device 100B applies
ultrasound pulses which generate echoes from the anterior and
posterior walls of an artery within the skin. Chemical sensing
using the electrochemical sensors 120B can begin with applying an
iontophoretic current from a positive terminal (e.g., anode+) to a
negative terminal (e.g., cathode -) that allows the
electro-repulsive delivery of a sweat stimulating molecule P+
(e.g., Pilocarpine nitrate). After pilocarpine delivery, stimulated
sweat containing biomarkers (e.g., including but not limited to
lactate, caffeine, and alcohol) is collected and quantified in the
left side of the sensor device 100B. The iontophoretic current
leads to osmotic flow of biomarkers (e.g., such as glucose) from
the interstitial fluid to the skin surface, allowing its collection
and analysis on the right side of the sensor device 100B.
[0082] FIG. 1D shows an illustration depicting various different
daily activities inputs performed by an individual and the
corresponding biomarkers (e.g., alcohol, caffeine, lactate, and
glucose) followed by the effect on the individual's system (body
response). The inputs are transduced and outputted as blood
pressure (BP), heart rate (HR), and electrochemical signals by the
device reflecting the body's physiological status.
[0083] FIG. 1E shows a schematic illustration depicting a
layer-by-layer layout of an example embodiment of a wearable
acoustic-chem sensor device 100, i.e., an acoustic-chem sensor
100E, showing a chemical sensing layer 120E, a substrate layer
125E, a physiological sensor layer 110E, and a substrate layer
115E. In this example, the substrate layer 125E and substrate layer
115E comprised styrene-ethylene-butylene-styrene block copolymer
(SEBS) material. For example, the use of SEBS or similar
styrene-based triblock copolymer, e.g., such as
styrene-isoprene-styrene block copolymer (SIS) or
styrene-butylene-styrene (SBS), are quite suitable for a flexible
substrate of the disclosed embodiments of the integrated
acoustic-electrochemical sensor devices, as these materials possess
a low melting point (e.g., .about.200.degree. C.), chemical
stability, biocompatibility, highly elastic mechanical behavior,
and solution processability. Such example properties of SEBS (and
SIS, SBS) grant softness, conformity to skin, heat-salability, as
well as compatibility to post-processing by solvent. Similarly, for
example, polymers such as polyurethane (PU) and polyethylene vinyl
acetate (PEVA) can be also consider suitable for such
applications
[0084] FIG. 1F shows two photos of the example acoustic-chem sensor
100E undergoing bending (panel (i)) and undergoing stretching
(panel (ii)).
[0085] FIG. 1G shows a diagram and corresponding data plot
depicting the detection mechanisms of the electrochemical sensors
120 for detection of two analytes, for some example embodiments of
the wearable acoustic-chem sensor device 100. FIG. 1G, panel i,
shows example amperometric measurements using enzyme-based sensors.
In this example (panel (i)), a Prussian blue (PB) working electrode
was modified with an analyte-associated enzyme, e.g., such as LOx,
GOx or AOx redox enzymes, allowing the biocatalytic oxidation of
lactate, glucose or alcohol molecules to pyruvate, gluconic acid or
acetaldehyde (product) respectively, along with the production of
hydrogen peroxide. In some example implementations, the
electrochemical reduction of the liberated hydrogen peroxide
(H.sub.2O.sub.2) to hydroxyl ions (OH--) can be performed in a
buffer, e.g., PBS pH 7.4, by applying a potential of -0.2 V. An
increase of negative current is observed by the increase in
concentration of chemical analyte (data plot of panel (i)). FIG.
1G, panel ii, shows example non-enzymatic measurements for
measuring an analyte such as caffeine. In this example (panel
(ii)), a carbon working electrode was modified with multi-walled
carbon nanotubes (MWCNTs) to facilitate a reaction at the sensor
surface for detection of caffeine. For example, during the sensing
process, caffeine is oxidized which results in the production of
uric acid analog molecules and electrons. In the example shown in
panel (ii), the carbon electrode modified with the MWCNT allowed
the pulse-voltammetric detection of caffeine following 30 s
accumulation at -1.2 V and scanning between +0.5 V and +1.5 V.
Other detection parameters shown in FIG. 1G include E.sub.step:
0.004 V; E.sub.pulse: 0.05 V; t.sub.pulse: 0.05 s; scan rate: 0.02
V/s. By increasing the concentration of caffeine, an increasing
oxidation signal is observed (data plot of panel (ii)).
[0086] FIG. 1H shows an illustration and data plot depicting an
example signal generation mechanism of the ultrasound transducer.
The pulsed ultrasound signal from the transducer is reflected from
the anterior and the posterior walls of the artery and collected by
the transducer. Signal processing of the ultrasound signal. The
time of flight (TOF) of the reflected echo can be converted into BP
via established transfer functions.
[0087] In the example embodiments of the device 100 shown in FIGS.
1A-1H, ultrasonic transducers are used for monitoring the BP and
HR, and electrochemical sensors are used for measuring the levels
of biomarkers. Through strategic material selection, layout design,
and fabrication innovation, rigid and soft sensor components were
integrated to create a wearable acoustic-chem sensor. For example,
in some embodiments, a customized array of piezoelectric lead
zirconate titanate (PZT) ultrasound transducers was integrated with
printed polymer composites via innovative solvent-soldering
process, into a single wearable conformal platform with high
mechanical resiliency and free of sensor crosstalk. Such rational
design overcomes engineering challenges related to the integration
of the different sensing modalities and materials to allow
real-time monitoring of cardiovascular parameters and biomarker
levels, in connection to parallel sampling of the interstitial
fluid (ISF) and sweat biofluids. The resulting epidermal hybrid
device can emit ultrasonic pulses and sense echoes from arteries,
while stimulating sweat and extracting ISF through iontophoresis
(IP), allowing simultaneous measurements of BP and HR, along with
multiple biomarkers in these biofluids.
[0088] As discussed further below, on-body trials were carried out
with multiple human subjects experiencing diverse activities and
stimuli (exercising, having alcohol, food, and caffeine, like those
illustrated in FIG. 1D). The correlations between metabolic
variations and hemodynamic activities, under these stimuli, were
monitored and evaluated. The improved sensor assembly process,
leveraging the Styrene-ethylene-butylene-styrene block copolymer
(SEBS)-based stretchable materials, allows the fast and reliable
fabrication of a stretchable and conformal epidermal sensor for
simultaneous acoustic and electrochemical sensing. Such a device
offers (i) comprehensive tracking of the effect of daily activities
and stimuli upon the users' physiological status, and (ii) enables
the collection of previously unavailable data towards understanding
of the body response to such stimuli, while addressing the critical
post-pandemic needs for remote telemetric patient monitoring.
[0089] In the example implementations of a wearable acoustic-chem
sensor device 100, SEBS was used as the stretchable and conformal
substrate to support the electrodes and connections printed with
customized inks (e.g., like that in FIG. 1E). The stretchable
substrate and inks allow the high conformity, flexibility (FIG. 1F,
panel (i)), and stretchability (FIG. 1F, panel (ii)) required for
wearable devices. The BP sensor includes an array of eight
piezoelectric transducers, which can be aligned with the carotid
artery upon applying on the neck to obtain optimal ultrasonic
signals. During sensing, the piezoelectric transducers were
activated with electrical pulses, transmitting ultrasound beams to
the artery, and the time of flight of the echoes from the anterior
and the posterior walls of the artery was analyzed to gauge the
dilation and contraction of arteries (e.g., like that illustrated
in FIG. 1C, 1H). The optimal BP signal can be selected from the
eight transducers with the best alignment to the artery and hence
the highest signal quality, thus ensuring reliable BP sensing
during movement where the patch may undergo some displacement.
Discussion regarding the fabrication of the example embodiment of
the wearable acoustic-chem sensor used in these example
implementations is discussed later in connection with FIG. 2M and
in Note 1. Example results of the BP sensor characterization is
discussed later in Note 3. In the example implementations, the
chemical sensing was realized through non-invasive sweat
stimulation (e.g., via transdermal pilocarpine delivery) at the IP
anode, alongside with ISF extraction at the IP cathode. For
example, lactate, alcohol and caffeine were monitored only in
sweat, while glucose was monitored only in ISF. Further information
regarding the simultaneous monitoring of ISF and sweat analytes via
iontophoresis is presented later in Note 4. Chronoamperometry (CA)
was used for electrochemical detection of the hydrogen peroxide
product of the glucose oxidase (GOx), lactate oxidase (LOx), and
alcohol oxidase (AOx) enzymatic reactions, while differential pulse
voltammetry (DPV) was used for the detection of caffeine. Example
embodiments of electrode modification and reaction mechanisms are
discussed in connection with FIG. 10. Example data showing the
analytical performance of each chemical sensor is shown in FIG. 1G,
as well as FIGS. 11-14, and Note 2.
[0090] FIG. 1I depicts another example embodiment of a wearable
acoustic-chem sensor device 1001 including the physiological
sensors 110 (e.g., including electrodes 111 and 113 operable to
measure ECG and blood pressure, respectively) and the
electrochemical sensors 120 (e.g., including lactate sensor 121 and
cortisol sensor 123 operable to measure lactate and cortisol
analytes, respectively). The physiological sensors 110 and
electrochemical sensors 120 are integrated onto a single flexible
substrate 1011. While the flexible wearable acoustic-chem sensor
device 1001 shows an example embodiment for some implementations,
yet many other combinations of physiological sensors 110 and
electrochemical sensors 120 can be integrated onto a single
substrate for embodiments of the wearable acoustic-chem sensor
device 100.
[0091] In some implementations of the example wearable
acoustic-chem sensor device 1001, the single flexible substrate
1011 is configured to a size that would span a length across a
user's heart, such that the ECG electrodes 111 are able to detect a
bioelectrical potential indicative of an electrocardiogram.
Whereas, in some embodiments, two separate wearable acoustic-chem
sensor device 1001 each comprising at least one ECG electrode 111
can be implemented by attaching the two separate wearable
acoustic-chem sensor device 1001 on opposite sides of the user's
heart, such that the two sensors 1001 are able to detect a
bioelectrical potential, via the respective ECG electrodes 111,
indicative of an electrocardiogram.
[0092] The single flexible substrate 1011 of the integrated
flexible biosensor 100 can be fabricated using a flexible
transparent substrate (e.g., ECOFLEX.RTM., PDMS, Polyurethane, and
so on) by using conductive screen-printed ink or by laser cutting
conductive material such as sheet metal. Example embodiments using
ECOFLEX can use a similar material, for example, as ECOFLEX
generally is made by a variety of polymers including PVC, ABS,
polyethylene, and polypropylene. The device 1001 may be designed
with the chemical sensor facing the skin, while the physical sensor
may be on the opposite side of the device.
[0093] FIGS. 1J and 1K show illustrative diagrams of an example
embodiments of a wearable integrated acoustic-electrochemical
sensor device 100, shown as wearable acoustic-chem sensor device
100J, attached to the skin of the body, e.g., on a user's neck. The
example wearable acoustic-chem sensor device 100J is configured to
simultaneously monitor the user's blood pressure and/or heart rate
via a physiological sensor 110J configured as an ultrasound sensor,
e.g., comprising an array of ultrasonic transducers 113J, in
parallel with a plurality of electrochemical sensors 120J to
non-invasively and electrochemically detect biomarker levels from
biofluids, such as sweat and/or interstitial fluid (ISF). The
example physiological sensor 110J and electrochemical sensors 120J
each are in electrical connection with electrical contact sites
107J via electrical interconnections 105J. The illustration of FIG.
1J and the exploded diagram of FIG. 1K both show an example
embodiment of the plurality of electrochemical sensors 120J that
includes a four-electrode contingent comprising four separate
detecting electrodes configured proximate to a reference electrode
and to a counter electrode, and with an iontophoretic electrode
spanning a relatively larger surface area at least partially around
the detecting electrode(s), counter electrode, and reference
electrode. In some embodiments of the wearable acoustic-chem sensor
device 100J, for example, the electrochemical sensors 120J that
include an iontophoretic electrode can include a hydrogel coupled
to the iontophoretic electrode, e.g., to aid in facilitating
extraction of an interstitial fluid from the skin and/or to aid in
inducing excretion of sweat from the skin, by triggered release of
sweat-inducing substances initially entrapped in the hydrogel
and/or by applied forces generated by electrical potentials applied
at the iontophoretic electrode. The example embodiment of the
plurality of electrochemical sensors 120J is discussed in further
detail (including but not limited to FIG. 39A), e.g., with respect
to example implementations for simultaneously monitoring four
analytes: glucose, lactate, caffeine, and alcohol. The illustration
of FIG. 1J and the exploded diagram of FIG. 1K both show the
example ultrasound sensor (of the physiological sensor 110J)
comprising the ultrasonic transducers 113J with a ground wire 114J
coupled to and spanning across the ultrasonic transducers 113 and
connecting an electrical ground interconnection 105J-G. In some
embodiments of the wearable acoustic-chem sensor device 100J, for
example, the physiological sensor 110J includes a hydrogel material
coupled to the ultrasonic transducers 113, e.g., to assist in
propagating acoustic signals generated at the ultrasonic
transducers 113 to the skin and propagating returned acoustic
echoes received from the skin to the ultrasonic transducers 113. In
some embodiments, for example, the electrical interconnection
structures 105J can be configured as serpentine interconnection
wires, which allow for stretching and bending on the flexible
substrate 101J.
[0094] FIG. 1L shows a block diagram of an example embodiment of an
electronic device 130 that can electrically interface to the
contact sites 107 of the wearable acoustic-chem sensor device 100
for various implementations. For example, the electronic device 130
can include an electrical circuit and/or a data processing unit to
process electrical signals as data received from the wearable
acoustic-chem sensor device 100. In various implementations, the
electronic device 130 is operable to store and execute software
applications to implement various sensing protocol algorithms
and/or implement various functionalities of the wearable
acoustic-chem sensor device 100. In various implementations, the
electronic device 130 can be implemented as a portable signal
processing and/or computing device, which can include a mobile
communications device, such as a smartphone, tablet or wearable
device, like a smartwatch, glasses, etc.; and/or, the electronic
device 130 can be implemented as a stationary signal processing
and/or computing device, such as a desktop computer and
amplifier.
[0095] In some embodiments, the electronic device 130 includes a
data processing unit 139 includes a processor 131 to process data,
a memory 132 in communication with the processor 131 to store data,
and an input/output unit (I/O) 133 to interface the processor 131
and/or memory 132 to other modules, units or devices, including
other external computing devices. For example, the processor 131
can include a central processing unit (CPU) and/or a
microcontroller unit (MCU) and/or a graphic processing unit (GPU).
For example, the memory 132 can include and store
processor-executable code, which when executed by the processor,
configures the data processing unit 139 to perform various
operations, e.g., such as receiving information, commands, and/or
data, processing information and data, and transmitting or
providing information/data to another device. In some
implementations, the data processing unit 139 can transmit raw or
processed data to a computer system or communication network
accessible via the Internet (referred to as `the cloud`) that
includes one or more remote computational processing devices (e.g.,
servers in the cloud). To support various functions of the data
processing unit 139, the memory 132 can store information and data,
such as instructions, software, values, images, and other data
processed or referenced by the processor. For example, various
types of Random Access Memory (RAM) devices, Read Only Memory (ROM)
devices, Flash Memory devices, and other suitable storage media can
be used to implement storage functions of the memory 132. In some
embodiments, the data processing unit 139 includes a wireless
communication unit 135, such as a wireless transmitter to transmit
stored and/or processed data or a wireless transceiver (Tx/Rx) to
transmit and receive data. The I/O 133 of the data processing unit
139 can interface the data processing unit 139 with the wireless
communications unit 135 to utilize various types of wired or
wireless interfaces compatible with typical data communication
standards, for example, which can be used in communications of the
data processing unit 139 with other devices, via a wireless
transmitter/receiver (Tx/Rx) unit, e.g., including, but not limited
to, Bluetooth, Bluetooth low energy, Zigbee, IEEE 802.11, Wireless
Local Area Network (WLAN), Wireless Personal Area Network (WPAN),
Wireless Wide Area Network (WWAN), WiMAX, IEEE 802.16 (Worldwide
Interoperability for Microwave Access (WiMAX)), 3G/4G/LTE/5G
cellular communication methods, NFC (Near Field Communication), and
parallel interfaces. In some embodiments, the data processing unit
139 includes a display unit 137, which can include a visual display
such as a display screen, an audio display such as a speaker, or
other type of display or combinations thereof. The I/O 133 of the
data processing unit 139 can also interface with other external
interfaces, sources of data storage, and/or visual or audio display
devices, etc. to retrieve and transfer data and information that
can be processed by the processor 131, stored in the memory 132, or
exhibited on an output unit (e.g., display unit 137) of the
electronic device 500 or an external device. For example, the
display unit 137 can be configured to be in data communication with
the data processing unit 139, e.g., via the I/O 133, to provide a
visual display, an audio display, and/or other sensory display that
produces the user interface of the software application. In some
examples, the display unit 137 can include various types of screen
displays, speakers, or printing interfaces, e.g., including but not
limited to, light emitting diode (LED), or liquid crystal display
(LCD) monitor or screen, cathode ray tube (CRT) as a visual
display; audio signal transducer apparatuses as an audio display;
and/or toner, liquid inkjet, solid ink, dye sublimation, inkless
(e.g., such as thermal or UV) printing apparatuses, etc.
[0096] Additional example embodiments including patch designs are
illustrated in FIGS. 39 and 40, which can be used for the
simultaneous monitoring of BP and multiple sweat-based chemical
markers (analytes), which are discussed later.
[0097] Example implementations were performed using some example
embodiments of the disclosed electrochemical-ultrasonic
transducer-based sensor technology, demonstrating simultaneously
detection and monitoring one or more bio-analyte markers and one or
more physiological markers.
[0098] Example Implementations of Wearable Acoustic-Chem Sensor
Device Embodiments
[0099] FIGS. 2A-2Q show diagrams, images and data plots depicting
example implementations characterizing example embodiments of a
multimodal wearable integrated acoustic-electrochemical sensor
device.
[0100] FIG. 2A shows an illustration of an example embodiment of
the wearable acoustic-chem sensor device 100, shown as wearable
acoustic-chem sensor device 200, highlighting the ISF analyte
sensor contingent and physiological sensor contingent for signal
interference crosstalk studies between the ISF electrochemical
sensor and the BP sensor. The example wearable acoustic-chem sensor
device 200 can include the one or more physiological sensors 110
comprising ultrasonic transducers, the one or more electrochemical
sensors 120, and the contact sites 107 disposed on the flexible
(bendable and/or stretchable) substrate 101, with and the
interconnection wires 105 disposed on and/or in the flexible
substrate 101, similar or the same as the example embodiment of the
device 100 shown in FIG. 1A. FIG. 2B shows a data plot depicting
the BP signal recording while applying and removing the CA
detection potential. FIG. 2C shows a data plot depicting the ISF
analyte electrochemical sensor signal recording while start and
pausing ultrasound signal generation with 30 s intervals within 3
min.
[0101] FIG. 2D shows an illustration of the example wearable
acoustic-chem sensor device 200, highlighting the sweat analyte
sensor contingent and physiological sensor contingent for signal
interference crosstalk studies between sweat electrochemical sensor
and the BP transducer. FIG. 2E shows a data plot depicting the BP
signal recording while applying and removing the CA detection
potential. FIG. 2F shows a data plot depicting the sweat analyte
electrochemical sensor signal recording while start and pausing
ultrasound signal generation with 30 s intervals within 3 min.
[0102] FIG. 2G shows photos of the example wearable acoustic-chem
sensor 200 under 20% vertical strain. FIG. 2H shows a plot
depicting envelopes of the raw echo signals before and after every
200 stretching cycles until 1000 cycles, and FIG. 2I shows a plot
depicting the electrochemical response every 200 stretching cycles
until 1000 cycles.
[0103] FIG. 2J shows photos of the example wearable acoustic-chem
sensor 200 under 20% horizontal strain (RSD=1.09%). FIG. 2K shows a
plot depicting envelopes of the raw echo signals before and after
every 200 stretching cycles until 1000 cycles, and FIG. 2L shows a
plot depicting the electrochemical response every 200 stretching
cycles until 1000.times. (RSD=12.24%).
[0104] FIG. 2M shows an illustrative diagram depicting a
fabrication method 250 of the example acoustic transducer of a
physiological sensor contingent, e.g., the piezoelectric lead
zirconate titanate (PZT) ultrasound transducer transfer process.
The method 250 includes a process 252 to deposit (e.g., drop cast)
an organic solvent 262 (e.g., toluene, xylene, benzene,
cyclopentane, n-pentane, cyclohexane, cyclohexanone, ethylbenzene,
acetone, methanol, ethanol, isopropanol, tetrahydrofuran, dimethyl
sulfoxide, or the mixture thereof) on the electrode pad 261 to
dissolve the material trace. The method 250 includes a process 254
to deposit a bonding material 264 (e.g., softened silver ink) for
bonding with an acoustic transducer structure 265, which includes
an acoustic transduction material, e.g., PZT, and to apply the
acoustic transducer structure 265 on the bonding material 264.
Other examples of an acoustic transduction material alternative to
or in addition to PZT include lead magnesium niobate-lead titanate
(PMN-PT) and/or polyvinylidene difluoride (PVDF). The method 250
includes a process 256 to deposit (e.g., drop cast) an organic
solvent (e.g., toluene) on transducers for bonding with a ground
wire 266. The method 250 includes a process 258 to binding the
ground layer to a reserved electrode channel.
[0105] FIG. 2N shows images of an example acoustic transducer
during an example implementation of the method 250, e.g., depicting
the adhesion of a PZT transducers to the substrate. Photos (i),
(ii), and (iii) of FIG. 2N shows images of the fabricated acoustic
transducer component under indentation, during horizontal
stretching, and after transferring the ground layer,
respectively.
[0106] FIGS. 2O-2Q show images depicting skin conformability and
mechanical integrity of an example wearable acoustic-chem sensor
device while twisting (FIG. 2O), bending (FIG. 2P), and after these
deformations (FIG. 2Q).
[0107] Crosstalk Study
[0108] Example implementations were performed to study potential
crosstalk in the example wearable acoustic-chem sensor device 200.
For example, the performance of an integrated sensor for
multiplexed simultaneous sensing requires reliable data generation
from the individual sensors, with no crosstalk between the two
sensing modalities. Here, the signal crosstalk between the acoustic
and electrochemical transducers of the wearable acoustic-chem
sensor device 200 was prevented by spatially separating both
components and using solid-state hydrogel materials for ultrasound
and sensing layers; also see FIGS. 31 and 32. For example, unlike
liquid gel, solid hydrogel layers will not smear at the interface
and thereby prevents potential circuit shorting between the
ultrasound transducer electrode and electrochemical sensing
electrodes, i.e., prevents signal crosstalk. For example, the
wearable acoustic-chem sensor 200 was designed with an optimal
distance between the individual detection compartments to ensure
successful acoustic BP and HR sensing, IP extraction, and
electrochemical monitoring. For example, the optimal distance
between the individual detection compartments can be configured in
a range of 5 mm to several centimeters, and more preferably in a
range between 100 mm to 1 cm or greater. In example implementations
discussed below, an optimal distance was determined to be 1 cm. The
distance can be determined based on the configuration of the
iontophoretic electrodes to ensure the net flow of interstitial
fluids to desired locations, for such embodiments that include one
or more iontophoretic electrodes. Depending on the size of the
sensor design and/or use of hydrogel material(s) for acoustic
coupling the acoustic sensor to skin and/or for ISF extraction or
sweat inducement, for example, the optimal distance between the
ultrasound transducer and the iontophoretic extraction electrode
can be above 0.1 cm, which can make sure no physical contact occurs
between the ultrasound hydrogel and the sensing hydrogel during
use.
[0109] As shown in FIG. 2A and FIG. 2D, the BP transducers were
located 1 cm below the chemical sensors, a distance optimized by
assessing the crosstalk between the neighboring sensors. In the
example implementations, the signal generation of the acoustic
sensor relies on high-voltage high-frequency pulses that may induce
signal drift in the chemical sensors, while the IP extraction,
potentiostatic sensing, and potential-sweep sensing may also induce
noises in the acoustic signals.
[0110] Possible crosstalk effect between the electrochemical and BP
sensors was evaluated by recording the corresponding signals during
on-body operations. The BP signals were acquired while the
potentiostatic electrochemical input was turned on and off
repeatedly to assess the effect of the electrochemical sensing of
the anodic sensor (e.g., see data plot of FIG. 2E and FIG. 29) and
cathodic sensor (e.g., see data plot of FIG. 2B) on the BP signal.
Similarly, the effect of the acoustic sensing on both sides of the
electrochemical sensing was examined by recording the amperometric
response while turning the acoustic pulses on and off repeatedly
every 30 s (FIG. 2C, 2F). Notably, for example, without
optimization as in the disclosed embodiments in accordance with the
present technology, the electrochemical detection was subject to
substantial signal interference due to the potential drift caused
by the biased voltage from the acoustic pulses, e.g., depicted in
data plots discussed later in connection with FIGS. 30 and 31.
[0111] Mechanical Performance
[0112] Example implementations were performed to study mechanical
performance of the example wearable acoustic-chem sensor device
200. For example, the mechanical stability is another crucial
factor that dictates the reliability of skin-worn sensors when
tensile deformations are expected. The impedance of the chemical
sensor and the contact resistance to the PZT transducers may vary
with the strain applied to the soft conformal device leading to
changes in the measured signals that affect the reliability of the
device. The stability of the PZT contact upon mechanical stress was
realized by developing a novel solvent-soldering method, i.e., the
method 250, which is illustrated in FIG. 2M, e.g., based on the
fast dissolution and room-temperature curing of SEBS-based
materials. During the assembly process, the PZT transducers can be
quickly mounted and bonded onto the SEBS substrate and connected to
the SEBS-based stretchable silver ink by wetting the electrode
surface with toluene. The solvent soldered PZT chips can thus be
securely bonded to the printed electrodes without delamination
during stretching deformations (as shown in FIG. 2N), with their
assembly efficiency largely improved.
[0113] The effects of stretching on the sensing performance were
assessed by stretching tests at 20% uniaxial strain. The example
device was stretched repeatedly along the vertical direction (FIG.
2G) and horizontal direction (FIG. 2J). The ultrasonic echo
signals, against a two-layered ECOFLEX, and the current (CA signal)
from the bare PB electrode (held at -0.2V), in buffer solution,
were recorded after every 200 cycles of stretching at 20% strain.
As shown in FIGS. 2H and 2K, although the intensity of the acoustic
transducer signal decreased slightly with stretching, the temporal
relationship between each peak that corresponded to two echoes did
not change, and hence the deformations did not affect the recorded
waveform. Similarly, the electrochemical sensors did not show
significant current change as the stretching cycle progressed
(FIGS. 2I, 2L, and FIGS. 17 and 18).
[0114] The example wearable acoustic-chem sensor device has also
shown good mechanical resilience after transferring it to the body.
FIGS. 20-2Q illustrate the twisting and bending of the sensor on
the skin. Mechanical resilience tests were performed also during
active 20% stretching deformation. The bare and enzyme-modified
electrochemical sensors were evaluated in vitro while under stress
in the horizontal and vertical directions, and the BP device was
used to capture the signal while turning the neck 90.degree.; the
glucose response did not change after or during the 100.times.
stretching (e.g., discussed later in connection with data plots of
FIGS. 20, 21, and 22). The SEM images depicting the surfaces'
structural changes of the printed stretchable silver and carbon
traces are displayed in FIG. 19, demonstrating that the printed
composites are not affected by the mechanical deformation.
[0115] Tracking Cardiovascular Activities and Biomarker Levels
[0116] Example implementations were performed to track
cardiovascular activities and biomarker levels using the example
wearable acoustic-chem sensor device 200. For example, the ability
of the wearable acoustic-chem sensor device to simultaneously
monitor dynamic cardiovascular parameters and biomarker
concentrations allows evaluating the effects of common daily
activities on an individual's physiological status and to
continuously collect data about their response to such everyday
activity. For example, the levels of lactate, glucose, alcohol, and
caffeine in our bodies can fluctuate due to common daily
activities, whose impact on our BP also varies based on an
individual's physical conditions. The simultaneous measurement of
biomarkers and BP allows the data collection of an individual's
responses to such daily activities. The device's ability to track
multiple biomarkers while capturing cardiac parameters can further
help deconvolute the additive effects of multiple stimuli on
physiological parameters, which holds significant implications
towards self-monitoring for personalized health management. In
order to study the effects of each activity upon the cardiac
parameters, measurements were performed before and after the
stimulus. Tests were performed by monitoring BP along with key
sweat and ISF biomarkers, corresponding to specific medical
situations.
[0117] Exercise, comprising any action which demands physical
efforts, has a major impact on the body's physiological response,
including changes in lactate levels, HR, and BP. During prolonged
exercising, blood and sweat lactate levels elevate due to metabolic
stress, HR increases to meet the muscle demand for oxygen, while BP
surges due to increased availability of vasodilatory mediators such
as nitric oxide. To study these effects, in the example
implementations, several volunteering subjects were asked to
perform stationary cycling at a fixed level for 30 min, followed by
20 min of resting. BP was recorded while the sweat was stimulated
before and after the exercise for the lactate measurements, and the
obtained BP and lactate level data were validated by a commercial
cuff-style blood pressure monitor and a blood lactate meter.
[0118] FIGS. 3A-3D show data plots depicting example data for an
on-body evaluation of the example hybrid acoustic-electrochemical
sensor device 200. FIG. 3A shows data associated with the BP/HR and
sweat lactate studies, including signal recording for BP/HR
performance before exercise (i) and after exercise (ii), bar
graphics represent the sensor validation using a commercial cuff
(white) and BP readings obtained with the ultrasound transducers
(green) (plot (iii)), signal recording for sweat lactate before
exercise (iv) and after exercise (v), and bar graphics represent
the sensor validation using a commercial blood lactate meter
(white) and readings obtained with the electrochemical sensor
(green) (in plot vi)). FIG. 3B shows data associated with the BP/HR
and sweat alcohol studies, including signal recording for BP/HR
performance before alcohol intake (i) and after alcohol intake
(ii), bar graphics represent the sensor validation using a
commercial cuff (white) and BP readings obtained with the
ultrasound transducers (blue) (plot (iii)), and signal recording
for sweat alcohol before alcohol intake (iv) and after alcohol
intake (v), bar graphics represent the sensor validation using a
commercial breathalyzer (white) and readings obtained with the
electrochemical sensor (blue) (plot (vi)). FIG. 3C shows data
associated with the BP/HR and ISF glucose studies. Signal recording
for BP/HR performance before food intake (i) and after food intake
(ii), bar graphics represent the sensor validation using a
commercial cuff (white) and BP readings obtained with the
ultrasound transducers (red) (plot (iii)), signal recording for ISF
glucose before food intake (iv) and after food intake (v), bar
graphics represent the sensor validation using a commercial blood
glucometer (white) and readings obtained with the electrochemical
sensor (red) (plot (vi)). FIG. 3D shows data associated with the
BP/HR and sweat caffeine studies, including signal recording for
BP/HR performance before caffeine intake (i) and after caffeine
intake (ii), bar graphics represent the sensor validation using a
commercial cuff (white) and BP readings obtained with the
ultrasound transducers (orange) (plot (iii)), signal recording for
sweat caffeine before caffeine intake (iv) and after caffeine
intake (v), bar graphics represent the sensor validation through
the standard addition method (white) and readings obtained with the
voltammetric sensor (orange) (plot (vi)).
[0119] As expected, significant changes in BP and HR were thus
observed in FIG. 3A after the exercise, increasing up to 150 mmHg
and 98 bpm, respectively (FIG. 3A at plots (i) and (ii)). Sweat
lactate also increased, with low lactate levels recorded in the
beginning and increased two-fold after the exercise (FIG. 3A at
plots (iv) and (iv)). The BP and sweat lactate data collected from
the device agreed well with the validation methods, as shown in
FIG. 3A at plots (iii) and (vi). It is worth noting that as no
exogenous drugs are used to affect the subject's BP, the relaxation
and contraction of the elastine- and collagen-rich central arteries
due to exercise can be considered negligible, and hence no
additional recalibration of the acoustic sensor was needed during
the experiment.
[0120] As another commonly seen unhealthy stimulus--excessive
alcohol consumption--has shown to increase cardiovascular risks via
alcohol-induced hypotension and hypertension. Alcohol may have
different effects on the BP, depending on the amount and frequency
of its consumption and genetic factors related to resistance or
sensitivity to alcohol. BP variations upon alcohol ingestion are
related to the direct vasodilation, surge in cortisol secretion,
and reduced insulin sensitivity. For the sensor experiments
focusing on alcohol as the stimulus, the BP and sweat alcohol level
were measured before and 20 min after drinking 200 mL of an
alcoholic beverage (19% vol.) (FIG. 3B). A commercial alcohol
breathalyzer was used for correlation with blood alcohol level. As
shown in FIG. 3B, the alcohol consumption resulted in an increased
HR (from 69 to 85 bpm) and BP (120 to 136 mmHg) of the volunteer
(FIG. 3B, plots (i) and (ii)). These results agree with studies
showing that a single alcohol intake by non-heavy drinkers can lead
to a temporary BP spike. It is worth noting that for heavy
drinkers, there might be a considerable BP morning surge that
greatly increases the risk of stroke. Simultaneously, the sensor
allows reliable detection of sweat alcohol, as this small polar
molecule can be found in sweat with a 1:1 correlation to blood
(FIG. 3B, plots (iv) and (v)).
[0121] Metabolites, such as glucose, can also affect the BP
waveform by changing the blood viscosity. Blood viscosity increases
under conditions of insulin resistance, altering the flux of blood
in the capillaries and hence the shape of the BP pulse. Studies
have shown that subjects with high blood pressure are prone to
significantly higher blood glucose levels. To test the effect of
the rise in glucose upon the BP, healthy non-diabetic subjects were
asked to consume a high sugary meal after fasting. The BP and ISF
glucose levels were recorded using the device before and 15 min
after consuming the food, with the glucose level validated using a
commercial glucometer at both times. As shown in FIG. 3C, plots (i)
and (ii), the sensor evaluation, during the food consumption
experiment, resulted in negligible changes in the BP and HR. In
contrast, the electrochemical biosensor readily detected changes in
the ISF glucose levels after the meal consumption (FIG. 3C, plots
(iv) and (v)). This data is within the expectation for the
non-diabetic subject, as glucose-induced BP changes occur only when
glucose levels increase significantly to alter the blood pumping
through the arteries, which is not common for non-diabetic
individuals whose glucose is readily regulated by the responsive
release of insulin.
[0122] Lastly, caffeine was chosen as another chemical stimulus
commonly used in many people's daily lives. Caffeine-intake is
known to lead to an increased BP through the inhibition of the
adenosine receptor and release of stress hormones, such as
norepinephrine or cortisol. These biochemical changes can result in
transient contractions of the arterial smooth muscle and influence
the vascular tone by phosphodiesterase inhibition. The effect of
caffeine on the BP varies, and are shown to be more pronounced in
hypertensive subjects. The epidermal BP/caffeine sensor patch was
evaluated on subjects with and without caffeine-intake habits, and
their BP and sweat caffeine were measured before and 30 min after
consuming a caffeine-rich (e.g., 80 mg) sugar-free energy drink.
The amount of caffeine in sweat was validated through a standard
additions voltammetric method, spiking caffeine to a collected
sweat sample (e.g., shown and discussed later in connection with
FIG. 24). As illustrated in FIG. 34, the on-body tests on a subject
with habitual caffeine-intake showed no significant changes in the
BP and HR after consumption of high caffeine doses, reflecting the
caffeine tolerance and healthy blood pressure levels of the
volunteer. In contrast, the BP variation was more pronounced for
the subject with no habitual caffeine intake, as shown in FIG. 3D.
The caffeine sensor displayed a flat DPV baseline response prior to
the caffeine intake, whereas the sweat DPV recorded 30 min after
the caffeine intake showed a distinct anodic peak current at 1.2 V,
corresponding to the caffeine oxidation (FIG. 3D, plots (iv) and
(v)). Current levels before and after the caffeine intake were
compared against the results obtained through the standard
additions method for caffeine, showing a good correlation between
both parameters (FIG. 3D, plot (vi)). Note that the in vitro
electrochemical characterization of the caffeine sensor in pH 4.5
showed current peaks around 1.1 V, e.g., shown and discussed later
in connection with FIG. 14. Such small potential shift reflects the
use of acetate-buffer loaded agarose gel over the caffeine sensor
for minimizing the effect of fluctuating sweat pH between
4.5-7.0.
[0123] Device Monitoring Multiple Stimuli
[0124] The example implementations included evaluation of the
example wearable acoustic-chem sensor device 200 in real-life
scenarios, where people usually experience multiple activities that
may have synergistic or counteracting effects on the body's
physiological response. The use of the example device 200 for
monitoring cardiovascular parameters along with multiple biomarker
levels was evaluated on subjects exposed to multiple stimuli. A
common example of counteracting effect to the glucose levels is
exercising along with food intake, as glucose can be quickly
consumed during exercise to produce energy. Exercise is also
expected to increase the BP and lactate levels in the subject, as
was shown in previous single-stimuli tests. To study this scenario,
the subject was asked to consume a sugar-rich meal, followed by
exercising on a stationary bike for 30 min, with the ISF glucose,
sweat lactate, and BP monitored before and after each step.
[0125] FIGS. 4A and 4B show data plots depicting example results
from an evaluation of a lactate, glucose, BP sensor performance and
an alcohol, glucose, BP sensor performance, respectively, using an
example embodiment of the wearable acoustic-chem sensor device 200.
FIG. 4A shows BP/HR signal recordings before exercise (plot (i))
and after exercise (plot (ii)), bar graphic comparison between BP
signal using a commercial cuff (white) and the ultrasound
transducer (green/red) (plot (iii)), electrochemical sensor signal
recordings for sweat lactate before (dotted line) and after (solid
line) exercising (plot (iv)), electrochemical sensor signal
recordings for glucose after having a meal and before exercising
(dotted line) and after exercise (solid line) (plot (v)), and a bar
graphic comparison between lactate levels in sweat using the
electrochemical sensor (green solid) and a commercial blood lactate
meter (green/white), glucose levels in ISF using the
electrochemical sensor (red solid) and blood using a blood glucose
meter (red/white) (plot (vi)). FIG. 4B shows BP/HR signal
recordings before (plot (i)) and after (plot (ii)) food and alcohol
intake, bar graphic comparison between BP signal using a commercial
cuff (white) and the ultrasound transducer (blue/red) (plot (iii)),
electrochemical sensor signal recordings for sweat alcohol before
alcohol intake (dotted line) and after alcohol intake (solid line)
alcohol intake (plot (iv)), electrochemical sensor signal
recordings for ISF glucose before food intake (dotted line) and
after food intake (solid line) (plot (v)), and a bar graphic
comparison between alcohol levels in sweat using the
electrochemical sensor (blue solid) and a commercial breathalyzer
(blue/white), glucose levels in ISF using the electrochemical
sensor (red solid) and blood using a blood glucose meter
(red/white) (plot (vi)).
[0126] As shown in FIG. 4A, normal systolic BP level, high glucose
levels (e.g., >100 mg/dL) and low lactate levels were observed
before the biking activity. After the exercise, glucose levels
decreased, accompanied by a considerable increase in the BP, HR,
and lactate level, as predicted from previous tests. Control
experiments--performed without any food or exercise--were used to
corroborate that the change in signal resulted solely from the
increase of lactate and glucose levels (e.g., shown and discussed
later in connection with FIG. 37). Overall, FIG. 4A illustrates
that the new sensor is able to capture the complex processes
resulting from the simultaneous food and exercise stimuli,
including the digestion of food to produce glucose as the energy
reservoir, the glycolysis reaction consuming the glucose and oxygen
to release energy, the increased BP and HR compensating for the
oxygen depletion, and the lactate generation during the hypoxic
condition in exercise.
[0127] The influence of the simultaneous intake of alcohol and
glucose on the BP and HR, simulating a typical alcohol consumption
during meals, was also studied on volunteering subjects. Based on
previous observations, increasing glucose levels are not expected
to cause significant changes in the BP of the subjects, whereas an
increasing BP is expected after the alcohol intake. Therefore, an
additive effect in the rise in BP and glucose is expected when
combining the intake of alcohol and sugary food. Moreover, the
digestion of alcoholic drinks, along with the reduced insulin
sensitivity caused by alcohol consumption, can further aggravate
the increase in the glucose level and BP. On the other hand,
excessive alcohol intake can lead to severe hypoglycemia and
hypotension, even when combined with glucose intake, particularly
for insulin-dependent diabetes subjects. Therefore, the
simultaneous monitoring of glucose and BP is important for
distinguishing the case of moderate or excessive drinking and
preventing drinking-induced accidents, especially for subjects with
underlying health conditions. Sweat alcohol, ISF glucose, and BP
signals were recorded in the fasting state, after the alcohol
consumption, and after the food intake. As shown on FIG. 4B, plots
(iv) and (vi), before any food or alcohol consumptions, blood
glucose and alcohol showed a typical non-diabetic fasting state
reading of 90 mg/dL glucose and a 0% BAC level, whereas increasing
BP, glucose, and alcohol signals were observed for 20 min after the
stimuli. The observed increase in BP following the alcohol intake
alone was 16 mmHg (FIG. 3B, plot (iii)), rising further to 20 mmHg
after the concurrent intake of sugary food (FIG. 4B, plot (iii)).
Such BP variations demonstrate the synergetic effect of combining
alcohol and glucose intakes on the BP. Smaller changes in HR were
observed following the alcohol and food intakes as compared to the
alcohol intake alone, indicating different mechanisms for the
increased BP. Control experiments, carried out without intakes of
food or drink, were used to corroborate that the observed signal
changes were solely due to the increase of alcohol and glucose
(e.g., shown and discussed later in connection with FIG. 38), as
supported by early findings.
[0128] Continuous BP and Biomarker Monitoring
[0129] The ability of the sensor to capture the dynamic biomarker
and BP fluctuations while performing physical activity was also
demonstrated. Physically active individuals are expected to have
lower resting BP, reducing considerably the risk of heart failure
events. The lower resting BP can further be reflected in a smaller
increase in BP during exercising, as physically active individuals
signal the body earlier to release nitric oxide (NO) to promote
enhanced vasodilation. Smaller increases in lactate levels are also
expected for active individuals compared with non-active ones. BP
is expected to decrease following intense exercise activity,
eventually returning to its original value, regardless of the
fitness level. Further, studies demonstrated a close relationship
between the magnitude of the post-exercise BP decrease and the
lactate levels, showing that elevated blood lactate levels after
high-intensity exercise promotes larger differences between pre-
and post-exercise BP values. Such complex dynamic processes thus
require the hybrid sensor to operate continuously for capturing
these real-time fluctuations throughout the activity. Subjects with
different fitness levels (physically active and non-active) were
asked to perform a 30 min cycling activity at constant intensity
while wearing the device (during the whole experiment), and their
BP and sweat lactate levels were monitored continuously until the
exercise was stopped. IP was not used for this portion of the
study, as sweat was generated spontaneously from the activity.
Validation data were also recorded before, 10 min into, and after
the exercise.
[0130] FIGS. 5A and 5B show data plots depicting example results of
an evaluation of continuous lactate, BP, HR sensor performance for
an actively fit volunteer and for a sedentary volunteer,
respectively, using an example embodiment of the wearable
acoustic-chem sensor device 200. FIG. 5A shows continuous
lactate/BP/HR performance for an actively fit volunteer, including
continuous signal recordings showing sweat lactate profile during
stationary biking (plot (i)), bar graphics showing validation using
a commercial blood lactate meter (white) and electrochemical sensor
readings (green) (plot (ii)), BP/HR signal recordings before
stationary biking (plot (iii)), during stationary biking (plot
(iv)), and after stationary biking (plot (v)), and a bar graphic
comparison between BP signal using a commercial cuff (white) and
the ultrasound before (green), during (red), and after (purple) of
the exercise performance (plot (vi)). FIG. 5B shows continuous
lactate/BP/HR performance for sedentary volunteer, continuous
current recording showing sweat lactate profile during stationary
biking (plot (i)), bar graphics showing validation using a
commercial blood lactate meter (white) and electrochemical sensor
readings (green), (plot (ii)), BP/HR signal recording before
stationary biking (plot (iii)), during stationary biking (plot
(iv)), and after stationary biking (plot (v)), and a bar graphic
comparison between the BP signal of a commercial cuff (white) and
the ultrasound before (green), during (red), and after (purple) of
the exercise activity (plot (vi)).
[0131] As shown in FIG. 5A (for the physically active subject) and
FIG. 5B (for the sedentary subject), a considerably higher sweat
lactate level and increased BP values were observed during the
exercise for the sedentary subject compared to the active subject.
Higher HR, BP, and sweat lactate levels are expected during
exercise for the non-active subjects due to the elevated
catecholamine levels compared to physically active subjects,
leading to differences in BP depending on the fitness levels and
cardiovascular system. To address the potential effect of sweat
electrolytes upon the activity of the Prussian Blue transducer, a
sufficiently negative applied potential (-0.2 V) was used, which
accommodates small possible shifts in the PB peak potential (e.g.,
shown in a data plot of FIG. 43). An example embodiment of an
advanced patch design, shown in FIG. 40, demonstrated it was able
to perform parallel potentiometric measurements of sweat
electrolyte levels for correcting the electrolyte effect. Example
on-body data is shown in FIG. 44, which is discussed later.
[0132] The example implementations described herein demonstrate the
first example of a conformal skin-worn device capable of
simultaneous monitoring of BP, HR, and multiple biomarkers. This
advance has been realized by elegantly addressing major engineering
challenges in integrating rigid ultrasound transducers and soft and
stretchable electrochemical sensors into a single flexible and
stretchable platform while ensuring mechanical performance and
avoiding signal crosstalk. The example SEBS-based solvent-soldering
process has greatly simplified the assembly of a sensor with
complex structure while ensuring reliable mechanical behavior and
continuous epidermal BP and biomarker signal recordings under
different chemical and physical stimuli and activities. Signal
crosstalk between the acoustic and electrochemical transducers was
prevented by spatially separating both components and using
solid-state sensing hydrogel layers. Repeated mechanical
deformation tests demonstrated outstanding durability and
reliability of the electrochemical and acoustic sensors.
[0133] Such simultaneous acoustic and electrochemical sensing
offers continuous monitoring of the users' physiological status and
its response to multiple everyday activities and stimuli. This
example multimodal wearable electrochemical/acoustic-physiological
sensing platform has thus been shown useful for correlating common
daily activities, such as exercise, drinking, and eating, with
changes in BP, HR, and biomarker levels. The example results
support the possibility of developing more advanced hybrid wearable
sensors that involve complex integration of chemical and physical
sensors on a single conformal platform for simultaneously
monitoring multiple relevant parameters. Such sophisticated
integration of reliable and comprehensive epidermal sensors was
realized with the judicious material selection, optimized
structural engineering, and novel high-throughput fabrication
process.
[0134] The disclosed wearable, integrated acoustic-electrochemical
sensor devices can be fully integrated in a miniaturized
electronics package, with integrated ultrasound and
multi-potentiostatic capabilities, along with signal processing and
wireless transmission functionalities. For example, the integrated
acoustic-electrochemical sensor device can interface with an
electronic device to provide multiplexed sensing modalities,
wireless communications, and display in a singular, wearable
platform. The wearable, integrated acoustic-electrochemical sensor
device can include a standalone acoustic sensing interface circuit,
coupled with artificial intelligence-aided signal processing.
[0135] It is envisioned that the disclosed wearable, integrated
acoustic-electrochemical sensor device can be included in a fully
integrated multiplexed wearable health monitoring device that
offers significant new insights into the health and physiological
status of individuals towards the prevention and management of
chronic diseases. The disclosed wearable, integrated
acoustic-electrochemical sensor device represents an important step
towards multimodal wearable sensors that fuse acoustic and
electrochemical sensors towards more comprehensive monitoring of
human physiology and a successful telehealth transformation. The
wearable, integrated acoustic-electrochemical sensor device is
envisioned to pave the way into a new field of skin-conformal tools
capable of providing important, high-quality, and high-density
information regarding the status of human health, and lays the
foundation for next-generation wearable patches capable of hybrid
chemical-electrophysiological-physical monitoring.
[0136] Example Methods of Fabrication and Implementations
[0137] Materials and Reagents
[0138] Example materials and reagents used in the example
implementations of the wearable acoustic-chem sensor device 200,
described herein, include chitosan, acetic acid, bovine serum
albumin (BSA), L-lactic acid, sodium phosphate monobasic
(NaH.sub.2PO.sub.4), sodium phosphate dibasic (Na.sub.2HPO.sub.4),
D(+)-glucose, glucose oxidase (GOx) from Aspergillus niger type X-S
(EC 1.1.3.4), Nafion.RTM., agarose, pilocarpine nitrate, Prussian
blue (soluble), toluene, ethanol, and silver flakes were obtained
from Sigma-Aldrich (St. Louis, Mo.). Graphite powder was purchased
from Acros Organics (USA). Lactate oxidase (LOx) (activity 101 U
mg.sup.1) was purchased from Toyobo Corp. (Osaka, Japan). SEBS
(G1645) was received from Kraton Corporation (Houston Tex., USA)
while ECOFLEX.RTM. 00-30 was purchased from Smooth-on Inc. (Easton
Pa., USA). Super-P carbon black was obtained from MTI Corporation
(Richmond, Calif., USA). The ultrasound gel pad (AQUAFLEX.RTM.) was
purchased from Parker Laboratories Inc. (Fairfield, N.J., USA).
Reagents were used without further purification.
[0139] Sensor Fabrication, Assembly and Electrode Modification
[0140] For the example implementations, screen-printing was carried
out using a semi-automatic MPM-SPM printer (Speedline Technologies,
Franklin, Mass.) and custom stainless-steel stencils developed
using AutoCAD software (Autodesk, San Rafael, Calif.) and produced
by Metal Etch Services (San Marcos, Calif.), with dimensions of 12
in .times.12 in and 125 .mu.m thickness. The electrodes were
printed layer-by-layer, as illustrated in FIG. 6. Bulk PZT was used
for ultrasound transducers, which were diced (Disco Automatic
Dicing Saw DAD3220) into 0.8 mm by 3 mm rectangular-shaped pixels
and sandwiched by two layers of stretchable silver inks as
electrodes. The connection between the transducers and the silver
traces was realized by adding a toluene droplet to the printed
silver traces and placing the transducers on the softened ink.
After attaching the PZT transducers, the screen-printed ground
connection was placed on the sensor by dissolving the printed
traces in a similar fashion. Details of the transducer assembly
process are illustrated in FIG. 7. The biosensor electrodes were
subsequently modified by drop casting the respective enzymes and
polymer layers. Details of the ink formulation, printing and
assembling processes, and the individual drop casting protocols for
different biosensors are discussed later in Note 1.
[0141] Sensor In Vitro Calibration
[0142] For the example implementations, fabricated sensors,
including the lactate, glucose, alcohol and caffeine biosensors,
and the PZT acoustic sensors, were calibrated separately in in
vitro settings. The biosensors were calibrated by using 0.1 M PBS
(pH 7.4) or 0.01 M acetate buffer (pH 4.5) with successive spiking
of corresponding analytes, and recording the corresponding CA (for
lactate, glucose and alcohol) and DPV (for caffeine). Protocols of
the example in vitro biosensor calibrations are discussed in detail
in Note 2. The calibration of the BP waveform is discussed in Note
3.
[0143] Sensor Mechanical Tests
[0144] For the example implementations, mechanical testing was
conducted via controlled stretching tests. A programmable motorized
linear stage (X-LRQ, Zaber Technologies Inc.) was used for
stretching the device with controlled strain and speed. One of the
edges of the printed device was taped at the fixed end of the stage
and the other to the moving end of the stage. The device was
firstly stretched at 3 mm/s speed to 120% of its original length in
the horizontal direction and release back to its original size at
the same speed. This process was programmed to be repeated 200
times, so that the device could be taken from the stage for
measurements before remounting back for subsequent stretching. The
process was repeated until 1,000 cycles of stretching were
completed (FIGS. 17 and 18). This process was repeated for
deformations on the vertical directions using the same device.
Electrochemical tests under constant deformation of 120% stretching
(horizontal and vertical directions) were conducted after every 200
stretching cycles, up to 1000 stretching deformations (FIGS. 20 and
21). The resiliency of the sensor was inspected visually by
attaching the device to the skin and subjecting it to various
deformations (FIGS. 2G, 2J); the corresponding printed surfaces,
before and after repeated stretching, were also characterized via
SEM imaging (e.g., FEI Quanta 250), as shown in FIG. 19.
[0145] Sensor Crosstalk Tests
[0146] For the example implementations, crosstalk between the
acoustic and electrochemical signals was analyzed on-body by
monitoring the changes in one signal while the other signal was
generated intermittently. For analyzing the co-sensor interference
from the CA electrochemical measurement to the BP waveform, the BP
signals were recorded continuously for at least 4 s in two stages.
Initially when the CA was already being performed by applying a
potential of -0.2 V to the electrochemical sensors, and when the
detection potential was turned on after the BP recording had
already started. The interference tests from the BP sensor on the
CA measurements were performed in the same fashion for the anodic
and cathodic sensors as follows. For analyzing the crosstalk effect
of the acoustic signal generation upon the electrochemical signal
acquisition, the CA signal was recorded continuously for 180 s
while the electric pulses for the BP measurements were delivered to
the PZT transducer in an off-on-off-on-off-on pattern with a period
of 30 s for each phase. The crosstalk from the differential pulse
voltammetry (DPV) to the acoustic signal was evaluated in the same
fashion as in the CA tests as follows. The effect of the acoustic
signal upon the caffeine sensor was evaluated in two stages, first
by recording the DPV signal while the BP recording was being
applied, following by terminating the BP signal when the DPV
reached the peak potential and by recording the DPV signal prior to
initiating the BP acquisition at peak potential. The corresponding
data obtained are included in FIGS. 2A-2Q and FIG. 29. Signal
generation and data acquisition were performed using .mu.Autolab
III electrochemical analyzer (Metrohm) for the chemical sensors and
the 5077PR pulser-receiver (Olympus) for the acoustic sensors. The
potentiostat was configurated with +/-5 V voltage and 1 mA current
limit to avoid overcurrent/overvoltage. The device was inspected
visually to ensure that the transducers were fully covered by the
SEBS substrates for insulation. No capacitive coupling,
short-circuiting nor breakdown conduction were observed during the
experiment.
[0147] Sensor On-Body Test Protocols
[0148] Epidermal evaluation of the device was performed on healthy
consenting subjects with no prior history of heart conditions,
diabetes, or chronic pain, and in strict compliance with the
protocol approved by the Institutional review board (IRB) at the
University of California, San Diego. The example device was placed
on the neck of the volunteers for all on-body evaluations. Prior to
every set of measurements using the integrated sensor, the glucose,
lactate, alcohol, and BP signals were validated with a commercial
glucometer (ACCU-CHEK, USA), blood lactate (NOVA biomedical, USA),
breathalyzer (BACtrack S80 Pro) and FDA approved blood pressure
cuff (LOVIA, USA), respectively. Caffeine concentrations were
estimated by standard addition methodology using the collected
sweat (e.g., shown in FIG. 24). Sweat stimulation and ISF
extraction were realized simultaneously by using a .mu.Autolab III
electrochemical analyzer to apply a current density of 0.3 mA
cm.sup.-2 between the cathode and anode electrodes for 10 min.
Prior to sweat generation, a pre-conditioning step was carried out
on the skin by applying the same current density using agarose gels
in the cathode and anode compartments for 10 min, following by
immediate placement of the device with pilocarpine delivery gel on
the conditioned area. Before placing the sensor, the skin was
thoroughly cleaned with soap and alcohol wipes. The patch was
transferred to the skin by using a double-sided clean laser tattoo
transfer adhesive (Papilio, TM). Openings were made in the adhesive
film to expose the sensors and IP electrodes to the skin. For the
measurements, a single device was used for each volunteer to
perform the "before" and "after" tests. The device was kept in the
volunteer's neck during the entire experiment, unless otherwise
specified. Detailed of the hydrogels fabrication methods and skin
transfer processes are illustrated in FIGS. 26, 27, and 28.
[0149] The example on-body results were acquired using a benchtop
CHI 1230A electrochemical analyzer for the biosensors and 5077PR
pulser-receiver (Olympus) for the acoustic sensors. Food intake
refers to the intake of sugar-rich food (100 g cheesecake, 350
kcal, 22 g sugar). Alcohol intake refers to the intake of alcohol
(200 mL wine, alcohol 19% vol.). Caffeine intake refers to the
intake of a sugar-free caffeinated drink (248 mL, 80 mg caffeine).
Exercise refers to a 30-min exercise session on a stationary bike
with constant intensity followed by a 5-min cooling period.
[0150] Exercising: BP and lactate signals were acquired before and
after exercising for three healthy volunteers. The device was
removed from their skin during the 30 min stationary bike exercise
and kept in a wet chamber at room temperature. After exercising,
and following a 5 min cooling period, the volunteer's neck was
cleaned with soap and alcohol pads for replacing the same sensor in
the same area. The optimal BP signal after the sensor replacement
was selected by testing the PZT array in the BP sensor. The
influence of the example device removal/replacement on the signal
was studied, as illustrated in FIG. 16.
[0151] Alcohol Intake: BP and alcohol levels were measured before
and 20 minutes after the alcohol consumption. The device was kept
on the volunteer's neck during the entire experiment.
[0152] Food Intake: BP and ISF glucose signals were acquired in the
fasting state (16 hours) for three healthy volunteers and 15 min
after consuming the sugar-rich food. The device was kept on the
volunteer's neck during the whole experiment.
[0153] Caffeine Intake: Subject volunteer's caffeine levels in
sweat were monitored before and 30 min after consuming the
sugar-free caffeine drink. The device was kept on the volunteer's
neck during the entire experiment. For the on-body tests an agarose
gel loaded with acetate buffer pH 4.5 was used, covering only the
caffeine sensor. Prior to the caffeine ingestion, stimulated sweat
was collected for the standard addition caffeine determination.
[0154] Simultaneous Alcohol and Food Intake: The dual modality of
the sensor was tested by combining alcohol and food intakes. BP,
ISF glucose, and sweat alcohol levels were measured before and
after 20 min of the simultaneous consumption of an alcoholic
beverage and the sugar-rich food. The device was kept on the
volunteer's neck during the whole experiment.
[0155] Food Intake and Exercising: The dual modality of the sensor
was tested toward the monitoring of blood pressure, glucose, and
lactate levels. The subject was first asked to consume a sugar-rich
food. Fifteen minutes after the food consumption, ISF glucose,
sweat lactate and BP were measured. Next, the device was removed
from the subject, kept in a wet chamber at room temperature, and
the volunteer was asked to perform the physical exercise on a
stationary bicycle for 30 min followed by cooldown for 5 min. After
the cooldown interval the subject's skin was cleaned and the same
sensor was used for subsequent measurement of the SIF glucose,
sweat lactate, and BP levels.
[0156] Continuous Lactate and BP Sensing During Exercising: The
sensor was further tested by monitoring dynamic changes in BP and
sweat lactate while performing continuous physical activity.
Subjects with different fitness levels (physically active and
non-active) were asked to perform 30 minutes of cycling activity at
constant intensity while wearing the sensor. Iontophoresis and the
iontophoretic gels were not used for this portion of the study, as
sweat was generated spontaneously from the activity. The BP and
blood lactate were measured right before the start of the exercise
and the initial sweat lactate level was measured 5 min after
starting the exercise when sweat was firstly generated. Within
.about.10 minutes, BP and blood lactate signal were recorded again.
The BP and blood lactate were recorded also upon completion of
exercising for validation.
[0157] For some example embodiments of the wearable acoustic-chemo
sensor device 100, some examples of ultrasound transducer
structures in accordance the disclosed technology can include
features of flexible ultrasound transducers like those that
described in U.S. Patent Publication No. 2019/0328354 A1, titled
"Stretchable Ultrasonic Transducer Devices," which is incorporated
by reference in its entirety as part of the disclosure of this
patent document. For some example embodiments of the wearable
acoustic-chemo sensor device 100, some examples of electrochemical
sensors in accordance the disclosed technology can include features
of sensors described in U.S. Pat. No. 9,820,692 B2 and/or U.S.
Patent Publication No. 2018/0220967 A1, each titled "Wearable
Electrochemical Sensors," and U.S. Patent Publication No.
2017/0325724 A1 entitled "Non-Invasive and Wearable Chemical
Sensors and Biosensors," which are incorporated by reference in
their entirety as part of the disclosure of this patent
document.
[0158] Additional designs, data, and discussion of example
implementations and example embodiments of wearable, integrated
acoustic-electrochemical sensor devices, in accordance with the
present technology, are described below.
[0159] Note 1. Example Sensor Fabrication Protocols
[0160] Fabrication of the Styrene-Ethylene-Butylene-Styrene Block
Copolymer (SEBS) Substrate
[0161] A viscous SEBS resin was prepared by dispersing SEBS beads
in toluene with a weight ratio of 1:2. The mixture was left on a
linear shaker (Scilogex, SK-L180-E) overnight or until the mixture
became transparent and homogeneous. A PET film with non-stick
coating was used as the temporary casting substrate, and a doctor
blade set at 1 mm height was used to cast the SEBS resin into a
sheet on the PET substrate. The cast resin was firstly dried in
ambient environment for 1 h, followed by curing in a conventional
oven at 80.degree. C. for additional 1 h to remove the excess
solvent. The transparent, uniform SEBS film was peeled off from the
PET substrate for subsequent sensor fabrication.
[0162] Synthesis of the Stretchable Silver and PB Ink
[0163] The stretchable silver ink was synthesized by mixing silver
flakes, toluene, and SEBS, in a weight ratio of 4:2.37:0.63, in a
dual asymmetric centrifugal mixer (Flacktek Speedmixer, DAC 150.1
KV-K) with a speed of 1800 RPM for 10 min or until obtaining a
homogeneous ink. The stretchable PB ink was synthesized by mixing
super-P carbon black, graphite powder, PB, SEBS and toluene, in a
weight ratio of 0.5:3:1:1.26:4.74, in the mixer at 2150 RPM for 10
min or until the ink was homogeneous. Before printing the
stretchable inks, the ink viscosity was analyzed visually and, if
necessary, (.about.200 .mu.L) toluene was added and the ink was
centrifuged before use.
[0164] Printing of the Stretchable Electrodes
[0165] The prepared SEBS sheet was used as the stretchable
substrate for the printed electrodes. An example embodiment of a
fabrication method, e.g., providing a step-by-step printing
technique of sensor structures, for example embodiments of the
wearable acoustic-chemical sensor device 200 is illustrated in FIG.
6. An example implementation of the method 600 included, firstly,
using stretchable silver ink for printing the interconnections, the
iontophoresis (IP) electrodes, and the reference electrodes, on the
front part of the SEBS substrate. Next, the stretchable PB ink was
used to print the working and counter electrodes of the biosensors.
The SEBS substrate was then turned over and the interconnections
for the transducer array were printed on the backside of the SEBS
substrate (opposite to the printed PB ink), with one extra channel
reserved for connecting to the common ground of the transducers. A
separate piece of SEBS sheet was used to print the ground wire for
the transducers. The printed inks were cured in a conventional oven
at 80.degree. C. for 10 min after each printing step. Before using
the complete printed device, the stretchable silver reference
electrodes were treated by adding a droplet (10 .mu.L) of 0.1 M
FeCl.sub.3 in 0.1 M KCl on the printed surface for 20 seconds to
produce the Ag/AgCl layer.
[0166] Assembly of the PZT Transducers
[0167] The diced PZT transducers can be "solvent soldered" onto the
printed current collectors by firstly dissolving the junction
regions of the interconnections temporarily with a small volume of
toluene (.about.1 .mu.L), followed by placing the transducers onto
the softened trances to physically bond with the composite ink.
After placing the transducers onto the wetted interconnections, the
assembled device was left drying for 2 minutes in ambient
temperature. Afterward, the printed ground wire was carefully
aligned with the transducers and solvent soldered to the array, by
adding a droplet of toluene on each PZT transducer, in a similar
fashion. Lastly, the printed ground was "solvent soldered" to the
reserved channel of the interconnection array. Details of the
example fabrication method is illustrated in FIG. 7.
[0168] Biosensor Working Electrode Modification
[0169] PBS used in the electrode modification was prepared in 0.1 M
with a pH of 7.4. BSA solution was prepared with a concentration of
10 mg/mL in PBS. The chitosan solution was prepared by dissolving
chitosan in 0.1 M acetic acid with a concentration of 0.5 wt %. For
preparing the lactate biosensor, the chitosan solution was mixed
with LOx (40 mg/mL) in BSA solution, in a ratio of 1:1 (v/v),
followed by drop casting a 2 .mu.L aliquot of the mixture onto the
working electrode surface. For preparing the glucose biosensor, GOx
(40 mg/mL) in BSA solution, glutaraldehyde in water (5 wt %) and
Nafion in water (0.5 wt %) were mixed in a ratio of 1:1:0.33
(v/v/v), and a 1.5 .mu.L aliquot was drop cast onto the working
electrode surface. For preparing the alcohol biosensor, AOx (10-40
units/mg), BSA solution and the chitosan solution were mixed in a
ratio of 8:1:1 (v/v/v), and a 4 .mu.L aliquot of the mixture was
drop cast to the working electrode surface. After drying at room
temperature for 1 hr, 2 .mu.L of the chitosan solution was drop
cast to all previously enzyme-modified surfaces. Upon completing
the corresponding modification steps, the resulting biosensors were
stored at 4.degree. C. overnight before using. For the caffeine
sensor, a solution of 0.1 mg/mL of MWCNT was dispersed in 50% EtOH
(v/v in DI water) in an ultrasound bath for 10 min, and a 2 .mu.L
aliquot of the dispersed solution was drop cast onto the working
electrode surface. After drying at room temperature for 1 hr, 2
.mu.L of 0.01% v/v of Nafion in water solution was drop cast onto
the previously MWCNT-decorated working electrode and dried at room
temperature overnight. Schematic illustrations representing the
modified sensor components and corresponding electrochemical
reactions for these example embodiments are shown in FIG. 10.
[0170] Preparation of the IP Hydrogels
[0171] An ECOFLEX mold with multiple circular trenches with a
diameter of 18 mm and thickness of at least 1 mm was prepared for
forming the hydrogels. The anode hydrogel solution was prepared by
dissolving 120 mg agarose in 3 mL of 0.1 M phosphate-buffered
saline solution at 150.degree. C. under stirring until the agarose
was dissolved. The cathode hydrogel solution was prepared by
dissolving 120 mg agarose in 3 mL DI water. After agarose
dissolution, the temperature was immediately decreased to
60.degree. C. and 60 mg of pilocarpine was added under continuous
stirring. 300 .mu.L aliquots of the hot solutions (cathodic and
anodic gels) were added into each circular mold to allow them to
solidify. After the solution cooled down in the mold, the gels were
cut into the corresponding cathode and anode geometries and stored
in a wet chamber at 4.degree. C. before using. Details of the
example preparation method are illustrated in FIG. 26.
[0172] Note 2. Example Sensor In Vitro Characterization
[0173] PBS used in the example in vitro characterization of the
sensors is 0.1 M at pH 7.4, unless otherwise noted.
[0174] Lactate Sensor
[0175] The calibration curve of the lactate sensor was obtained
using an initial PBS droplet with 100 .mu.L volume on the sensor
surface. The solution was spiked with 1 .mu.L of 0.5 M lactate
solution to incrementally increase the concentration of the lactate
from 0 to 30 mM with CA at -0.2 V for 60 s after each spiking. The
selectivity of the lactate sensor was evaluated by performing CA
while spiking the PBS successively with lactate (2 mM), glucose
(0.2 mM), ascorbic acid (10 uric acid (60 and acetaminophen (10 The
stability of the lactate sensor was examined by performing 10
repetitive CA measurements of 2 mM lactate and calculating its
relative response changes in %. The example in vitro
characterization data of the lactate sensor is summarized in FIG.
11.
[0176] Glucose Sensor
[0177] The calibration curve of the glucose sensor was obtained
using an initial 100 .mu.L PBS droplet on the sensor surface. The
solution was spiked with 1 .mu.L of 0.1 M glucose solution to
incrementally increase the concentration of glucose from 0 to 10 mM
with CA at -0.2 V for 60 s after each spiking. The selectivity of
the glucose sensor was evaluated by performing CA while spiking the
PBS successively with glucose (2 mM), lactate (10 mM), ascorbic
acid (10 uric acid (10 .mu.M), and acetaminophen (10 .mu.M). The
stability of the glucose sensor was examined by performing 10
repetitive CA measurements of 2 mM glucose and calculating its
relative response changes in %. The example in vitro
characterization data of the glucose sensor is summarized in FIG.
12.
[0178] Alcohol Sensor
[0179] The calibration curve of the alcohol sensor was obtained
using an initial PBS droplet with 100 .mu.L volume on the sensor
surface. The solution was spiked with 1 .mu.L of 0.8 M ethanol
solution to incrementally increase the concentration of alcohol
from 0 to 32 mM with CA at -0.2 V for 60 s after each spiking. The
selectivity of the alcohol sensor was evaluated by performing CA
while spiking the PBS measured successively with ethanol (20 mM),
lactate (10 mM), glucose, (0.2 mM), ascorbic acid (10 .mu.M), uric
acid (60 .mu.M), and acetaminophen (10 .mu.M). The stability of the
alcohol sensor was evaluated by performing 10 repetitive CA
measurements of 2 mM ethanol and calculating its relative response
changes in %. The example in vitro characterization data of the
alcohol sensor is summarized in FIG. 13.
[0180] Caffeine Sensor
[0181] DPV was utilized for evaluating the caffeine sensor with the
following parameters: accumulation at -1.2 V for 30 s;
E.sub.initial: +0.5 V; E.sub.final: +1.5 V; E.sub.step: 0.004 V;
E.sub.pulse: 0.05 V; t.sub.pulse: 0.05 s; scan rate: 0.02 V/s. For
calibration curve tests, 100 .mu.L of 0.01 M acetate buffer (pH
4.5) were used to cover the sensing area of the device. The
background response was recorded repeatedly until the signal was
stable. The caffeine DPV response was then recorded after each
consecutive addition of 1 .mu.L of 1 mM caffeine in DI water for
obtaining 10 .mu.M increments of the caffeine concentration, up to
210 .mu.M. The selectivity of the caffeine sensor was evaluated by
performing DPV while spiking the acetate buffer successively with
caffeine (20 .mu.M), lactate (10 mM), ascorbic acid (10 .mu.M),
uric acid (60 .mu.M), and acetaminophen (10 .mu.M). The stability
of the caffeine sensor was examined by performing 10 repetitive DPV
measurements of 20 .mu.M caffeine and calculating its relative peak
current changes in %. The example in vitro characterization data of
the caffeine sensor is summarized in FIG. 14. For the example
on-body tests an agarose gel loaded with acetate buffer pH 4.5 was
used, covering only the caffeine sensor. A standard additions
voltammetric method, involving spiking caffeine to a collected
sweat sample, is used to validate the amount of caffeine in sweat.
This method was used once to construct the calibration plot for
correlating directly the wearable patch signal to sweat caffeine
concentrations.
[0182] Sodium Sensor
[0183] Open circuit potential (OCP) was used for evaluating the
sodium sensor. For the calibration curve, the sensor was incubated
for 30 minutes in an aqueous solution of 10 mM NaCl. After rinsing,
the response signals for 100 .mu.L of 0.1, 1, 10 and 100 mM NaCl
were recorded by consecutively replacing the solution on the
electrode surface. The selectivity of the sodium sensor was
evaluated by calibrating the sensor with different KCl
concentrations in a similar fashion as the NaCl calibration. The
stability of the sodium sensor was examined by recording the
response to 0.1 mM NaCl over 1 hour. The reversibility of the
sensor was performed by increasing and decreasing the NaCl
concentration on the sensor surface from 0.1 mM to 100 mM and back
to 0.1 mM.
[0184] Note 3. Calibration of Blood Pressure Waveform
[0185] The position of the artery walls was represented by the
flight time of the echo signals. By continuously tracking the echo
shift of the vessel walls, the arterial distension waveform could
be recorded (e.g., shown in FIG. 15). Then, based on an established
model, the arterial distension waveform could be transferred to
blood pressure waveforms.
[0186] The arterial blood pressure waveform p(t) is calculated from
the vessel distension waveform d(t) as follows (in Equations (1)
and (2)):
A .function. ( t ) = .pi. .times. .times. d 2 .function. ( t ) 4 (
1 ) p .function. ( t ) = p d e .alpha. .function. ( A .function. (
t ) A d - 1 ) ( 2 ) ##EQU00001##
where A(t) is the cross-sectional area of the artery, and d(t) is
the diameter of the target artery. Here, the artery is assumed to
be rotationally symmetrical. p.sub.d is the diastolic pressure.
A.sub.d is the diastolic arterial cross-section, and a is the
rigidity coefficient.
[0187] .alpha. can be calculated by the following equation
(Equation (3)):
.alpha. = A d .times. ln .function. ( p s / p d ) A s - A d ( 3 )
##EQU00002##
where A.sub.s is the systolic arterial cross-section, and p.sub.s
is the systolic pressure. The p.sub.d and p.sub.s are measured by
the commercial blood pressure cuff from the brachial site. Using
the above equations and a brief calibration for .alpha., p.sub.d,
and p.sub.s the accurate pressure waveform p (t) can be
obtained.
[0188] It is noted that the human blood vessel is assumed to be
elastic with negligible viscoelasticity. For subjects with normal
local vascular conditions or with slight local atherosclerosis, the
diameter of the vessel does not lag behind the pressure
waveforms.
[0189] Note 4. Simultaneous Monitoring of ISF and Sweat Analytes
Via Iontophoresis
[0190] An iontophoretic system has been used for the simultaneous
ISF extraction (cathode) and sweat stimulation by pilocarpine
delivery (anode). The extraction and delivery operations are
performed at the same time based on two mechanisms, as described in
detail in previously published work. In brief, a low-intensity
electrical current is applied to the skin using two electrodes
(cathode and anode). Iontophoretic gels with different compositions
are located under each electrode. On the anode compartment, a
pilocarpine-loaded gel is used for stimulating the sweat.
Pilocarpine is delivered inside the skin by electrical repulsion as
the pilocarpine molecule is positively charged and a positive
current is applied at the anode. Then, localized sweat production
occurs only in the stimulated area where the pilocarpine drug was
delivered (anode). Since no sweat stimulant drug is present on the
cathode compartment, no sweat is produced under the cathode
electrode. The iontophoretic gel, located in the cathode
compartment, is loaded with PBS buffer and a negatively charged
current is applied to attract the positively charged ions from the
ISF under the skin to the outside. A flow of negatively charged
ions is also attracted to the skin surface in the anode
compartment; however, as the skin is naturally negatively charged,
there is a net flow of ISF toward the cathode, which carries all
small neutral molecules in the same direction. ISF glucose can thus
be detected in the cathodic compartment. Accordingly, the
electroosmotic convective flow is responsible for the ISF glucose
extraction exclusively on the cathode compartment while the target
sweat analyte is detected at the anode.
[0191] Note 5. SEM Analysis from Mechanical Deformation
[0192] For examining the morphological change of the printed
composite electrode before and after stretching, two sets of
samples for the SEM imaging were prepared, before and after the
1500 times of stretching. For example, the change in the apparent
surface morphology can be caused by several factors, including the
difference in the individual samples, the contrast, and brightness
of the SEM image being taken, as well as the most importantly, any
physical damage caused by the stretching deformation. A key element
to examine was if there was any cracking, peeling, or delamination,
on the surface of the electrode. As shown in FIG. 18, the printed
silver composite, mainly used as the reference electrode and the
interconnection for the electrochemical sensors and the PZT chips,
has no apparent cracking, peeling, or delamination from the
substrate. The Prussian blue-carbon ink, used for the working and
auxiliary electrodes of the electrochemical sensors, has shown
minor cracking on the surface of the electrode. Such behavior is
expected as the formulation of the carbon-PB ink includes a high
loading of small-size, highly porous materials, which made the
composite less stretchable compared to the silver-based ink.
However, due to the small size and high active surface area of the
working electrode, such minor morphological change should have a
negligible effect on the sensing results and upon the sensitivity.
To support this view, chronoamperometric data in FIGS. 2G-2L shows
that the current response of the electrochemical. sensor is not
affected by the repeated stretching. Thus, the electrochemical
performance of the sensor was not impaired by the mechanical
deformation. The acoustic transducers, based solely on the silver
ink, rely mainly on the temporal resolution of the signals instead
of its intensity, and was also not affected by the stretching
deformation. Furthermore, additional mechanical tests to measure
the example sensor's performance during deformation were
implemented (e.g., as shown in FIGS. 20 and 21). The example
results also indicate no hysteresis due to the applied strain on
the printed electrodes. Thus, the example sensor was able perform
normally within the designed level of deformation. As a wearable
epidermal sensor, for some examples, the usage of the example
sensor could be up to a week, or in other examples, the usage of
the example sensor could be up to a month; the example sensor was
tested over 1000 times of repeated deformation at the strain of
20%, which is presumed to be extreme and unlikely to occur in real
life. Compared to bending and twisting, stretching deformation
applies the most mechanical stress to the printed materials. Due to
the use of hydrogels, abrasion on the electrode is also less likely
to occur. The repeated stretching tests were thus used as the most
rigorous test for the durability of the sensor. Overall, the
aforementioned supporting data reflect the stable and durable
performance of the integrated sensor.
[0193] Note 6. Sodium Ion Selective Electrode
[0194] The sodium selective membrane cocktail composition included
1 mg sodium ionophore X, 0.55 mg Na-TFPB, 33 mg PVC, and 65.45 mg
DOS dissolved in 660 .mu.L of THF (Fisher Chemical). The cocktail
was thoroughly mixed to dissolve all the components. The reference
cocktail was prepared by dissolving 78.1 mg PVT (Quimidroga S.A.)
and 50 mg NaCl in 1 mL methanol. Next, a 3 .mu.L aliquot of the
sodium selective membrane cocktail was drop-casted onto the working
carbon electrode and the reference electrode was modified by 3
.mu.L aliquot of the reference cocktail, followed by 1 .mu.L of
polyurethane (Tecoflex.RTM. SG-80A) dissolved in THF (15% w/w). The
modified Na-sensors were left to dry overnight before use.
Chemicals were obtained from Sigma Aldrich (St. Louis, Mo.), except
when specified otherwise.
[0195] FIG. 6 shows a diagram illustrating a fabrication method 600
for a layer-by-layer printing and assembling of an integrated
sensor, in accordance with the present technology. At a process
610, the method 600 includes using stretchable silver and PB ink to
print the pattern over the SEBS substrate. At a process 620, the
method 600 includes printing the stretchable serpentine
interconnection using the silver ink. At a process 620, the method
600 includes printing of the iontophoresis (IP) electrodes and the
reference electrodes using the silver ink. At a process 630, the
method 600 includes printing the counter (curved) and working
electrodes (round) using the PB ink. At a process 640, the method
600 includes printing an insulating layer to define the working
electrode area and insulate the interconnections using the SEBS
resin. At a process 650, the method 600 includes flipping the SEBS
substrate backside up and printing the interconnects (e.g.,
serpentine interconnects) for the transducers and ground using the
silver ink. At a process 660, the method 600 includes using a
conductive ink solvent (e.g., silver ink solvent) as an adhesive to
bond the transducer chips at terminuses of the silver
interconnects. At a process 670, the method 600 includes using the
conductive ink solvent to as an adhesive to bond a ground wire
structure to the other side of the transducers and connect it to
the reserved ground interconnect. After the process 670 of the
method 600, the partially fabricated device is ready for sensor
modifications. The method 600 can optionally include a process 680
to flip the sensor and implement a sensor surface modification
method to tailor the sensors for sensing targeted analytes.
[0196] FIG. 7 shows a diagram illustrating a method 700 for
assembly of ultrasound transducers for some example embodiments of
the wearable acoustic-electrochemical sensor devices, in accordance
with the present technology. At a process 710, the method 600
includes depositing an organic solvent (e.g., dipping toluene
droplets) on the connection pad of the interconnects, e.g., using a
pipette to partially dissolve the silver traces. At a process 720,
the method 700 includes placing the transducer on the softened pad
for bonding. At a process 730, the method 700 includes repeating
the processes 710 and 720 and aligning all the transducers along
the pattern. At a process 740, the method 700 includes depositing
an organic solvent (e.g., dipping toluene droplets) on transducers
and the ground interconnect pad, e.g., using a pipette. At a
process 750, the method 700 includes applying the ground wire to
the transducers and connecting to the ground interconnect pad.
[0197] FIG. 8 shows a diagram and an image depicting transducer
dimensions and conformability. Diagram (a) of FIG. 8 shows
dimensions of example PZT transducer pixels and a transducer array
of the PZT transducers, including one example of the size and
spacing of the transducer pixels in the array. For example, eight
PZT transducer pixels are configured to have a height of about 250
.mu.m, a width of about 1 mm, and a length (or depth) of about 4
mm; and each of these PZT transducer pixels are spaced apart by
about 200 The aspect ratio of each pixel is controlled to be
smaller than 0.3 (w, l>3 h) to ensure that the PZT vibration is
in thickness mode with accurate frequency range or particular
frequency. For example, the aspect ratio of the ultrasound
transducer array pixels can ensure the acoustic transducer material
(e.g., PZT) vibration is in thickness mode in a frequency range of
2 MHz to 10 MHz, and preferably, for example, a frequency range of
5 MHz to 9 MHz. Also, for example, the aspect ratio of the
ultrasound transducer array pixels can ensure the acoustic
transducer material vibration is in thickness mode at a particular
frequency, such as 7 MHz. The ideal frequency or frequency range is
selected based on a balance between signal acquisition interests
and acquisition system complexities, e.g., as a higher frequency of
the acoustic signals may provide for better image quality, but the
higher the frequency requires more complexity in the acquisition
system (due to higher acquisition rates). Image (b) of FIG. 8 shows
a photo of the device on a spherical surface to demonstrate the
conformability of the fabricated transducer array.
[0198] FIG. 9 shows images depicting adhesion of the PZT
transducers to the substrate. Photos (a-c) are images of the
pristine device before any deformation. Image (a) shows, the device
during horizontal stretching. Image (b) shows the device under
indentation. Image (c) shows the PZT transducers remained well
attached to the printed silver traces after the deformations.
[0199] FIG. 10 shows diagrams illustrating example electrochemical
sensor electrode modifications and reaction mechanisms. The
examples shown in FIG. 10 only illustrate the modified working
electrode, but it is understood that the electrochemical sensor
also includes one or more additional electrodes, e.g., such as a
reference electrodes and/or a counter electrode, to perform the
described electrochemical reactions.
[0200] FIG. 10, diagram (a) illustrates an example lactate sensor
modification. The example lactate sensor includes a printed,
flexible (e.g., stretchable and/or bendable) carbon-based electrode
with embedded Prussian blue (PB) redox mediator probe 1003 for
hydrogen peroxide reduction. The PB-embedded electrode 1003 is
modified with lactate oxidase (LOx) enzymes 1002 immobilized with a
drop-cast polymer, e.g., chitosan stabilizer 1001. As shown in
diagram (a), a LOx reaction with lactate leads to the formation of
hydrogen peroxide and pyruvate. Further, a PB-based electrode
transducer transforms the hydrogen peroxide product to hydroxyl
ions (OH--) for selective lactate detection. For example, when
lactate is present at the lactate sensor, the LOx reaction leads to
the formation of hydrogen peroxide and pyruvate, which the hydrogen
peroxide is further reduced by the PB and its reductive current
detected at the carbon-based electrode. The current can be
therefore correlated with the lactate concentration upon
calibration.
[0201] FIG. 10, diagram (b) illustrates glucose sensor
modification. The example glucose sensor includes a printed,
flexible (e.g., stretchable and/or bendable) carbon-based electrode
with embedded Prussian blue (PB) redox mediator probe 1003 for
hydrogen peroxide reduction. The PB-embedded electrode 1003 is
modified with glucose oxide (GOx) enzymes 1004 cross-linked with
glutaraldehyde and immobilized with Nafion to form a
glutaraldehyde/Nafion layer 1005. As shown in diagram (b), the GOx
reaction with glucose leads to the formation of hydrogen peroxide
and gluconic acid. The PB-based electrode transducer offers
specific detection of the peroxide product towards selective
glucose detection. For example, when glucose is present at the
glucose sensor, the GOx reaction leads to the formation of hydrogen
peroxide and gluconic acid, which the hydrogen peroxide is further
reduced by the PB and its current detected at the carbon-based
electrode. The current can be therefore correlated with the glucose
concentration upon calibration.
[0202] FIG. 10, diagram (c) illustrates alcohol sensor
modification. The example alcohol sensor includes a printed,
flexible (e.g., stretchable and/or bendable) carbon-based electrode
with embedded Prussian blue (PB) redox mediator probe 1003 for
hydrogen peroxide reduction. The PB-embedded electrode 1003 is
modified with alcohol oxidase (AOx) enzymes 1006 immobilized with a
drop-cast polymer, e.g., chitosan stabilizer 1001. As shown in
diagram (c), the AOx reaction with its ethanol substrate results in
the formation of hydrogen peroxide and acetaldehyde. The PB-based
electrode transducer offers specific detection of the peroxide
product towards selective alcohol detection. For example, when
ethanol is present at the alcohol sensor, the AOx reaction leads to
the formation of hydrogen peroxide and acetaldehyde, which the
hydrogen peroxide can be further reduced by the PB and its
reductive current detected at the carbon-based electrode. The
current can be therefore correlated with the ethanol concentration
upon calibration.
[0203] FIG. 10, diagram (d) illustrates caffeine sensor
modification. The example caffeine sensor includes a
screen-printed, flexible (e.g., stretchable and/or bendable)
carbon-based electrode 1012 modified with carbon nanotubes (CNTs)
1014 immobilized with Nafion 1016, e.g., for increase effective
electrochemical active area. As shown in diagram (d), the anodic
oxidation of the caffeine analyte results in the production of uric
acid and electron flow. The DPV peak current corresponds to the
caffeine concentrations. For example, when caffeine is present at
the caffeine electrode, the sensor is scanned via differential
pulse voltammetry from 0.8-1.8 V and the oxidation peak current of
caffein is measured, corresponding to the oxidation of caffeine at
the applied high potential on the surface of the CNT electrodes.
The concentration of the caffeine can be correlated with the peak
current upon calibration.
[0204] FIG. 11 shows data plots depicting example in vitro
characterization data of an example lactate sensor. Plot (a) shows
the lactate sensor's amperometric response to successive additions
of 5 mM lactate from 0 to 30 mM. Plot (b) shows a lactate
calibration curve based on the data of Plot (a). Plot (c) shows the
evaluation of the lactate sensor selectivity in the presence of
lactate (LA, 2 mM), glucose (GLU, 0.2 mM), ascorbic acid (AA, 10
.mu.M), uric acid (UA, 60 .mu.M) and acetaminophen (AC, 10 .mu.M).
Plot (d) shows the stability of the lactate: 10 repetitive
measurements of 2 mM lactate.
[0205] FIG. 12 shows data plots depicting example in vitro
characterization data of an example glucose sensor. Plot (a) shows
the glucose sensor's amperometric response to successive 1 mM
glucose additions from 0 to 10 mM. Plot (b) shows the glucose
calibration curve based on the data of Plot (a). Plot (c) shows the
evaluation of the glucose sensor selectivity in the presence of GLU
(2 mM), LA (10 mM), AA (10 .mu.M), UA (10 .mu.M) and AC (10 .mu.M).
Plot (d) shows the stability of the glucose sensor: 10 repetitive
measurements of 2 mM glucose.
[0206] FIG. 13 shows data plots depicting example in vitro
characterization data of an example alcohol sensor. Plot (a) shows
the glucose sensor's amperometric response to successive 8 mM
ethanol increments from 0 to 80 mM. Plot (b) shows the calibration
curve of the alcohol sensor based on the data of Plot (a). Plot (c)
shows a selectivity test in the presence of AL (20 mM), GLU (0.2
mM), AA (10 .mu.M), LA (10 mM), UA (60 .mu.M) and AC (10 .mu.M).
Plot (d) shows the stability of the alcohol sensor: 10 repetitive
measurements of 20 mM alcohol.
[0207] FIG. 14 shows data plots depicting example in vitro
characterization data of an example caffeine sensor. Plot (a) shows
the caffeine sensor's DPV response of increasing caffeine additions
in 10 .mu.M steps from 0 to 200 .mu.M. Plot (b) shows a
corresponding caffeine calibration curve based on the data of Plot
(a). Plot (c) shows the evaluation of the caffeine sensor
selectivity in the presence of CF (20 .mu.M), GLU (0.2 mM), LA (10
mM), AA (10 .mu.M), UA (60 .mu.M) and AC (10 .mu.M). Plot (d) shows
the stability of the caffeine sensor: 10 repetitive DPV
measurements of 20 .mu.M caffeine.
[0208] FIG. 15 shows data plots depicting ultrasound transducer
characterization on a phantom. Plot (a) depicts a radio frequency
signal showing the anterior wall and posterior wall of the carotid
artery phantom. Plot (b) shows vessel wall displacement with
increased intravascular pressure on phantom. Plot (c) shows
periodic vessel distension induced by the inflator.
[0209] FIG. 16 shows a diagram and a data plot illustrating optimal
channel selection for accurate artery diameter tracking. Diagram
(a) shows the optimal channel was determined by calculating the
time of flight (ToF) of the ultrasound signal. Data plot (b) shows
the raw RF signal showing different ToF values from two adjacent
transducers. In the example implementations, only the maximum ToF
between anterior wall and posterior wall would be recognized as the
`diameter` of the artery. Thus, the accurate diameter tracking
could be guaranteed by the channel selection.
[0210] FIG. 17 shows data plots depicting electrode electrochemical
stability under repeated stretching tests. Plots (a) and (b) show
the CV response of the PB electrode every 200 cycles in a
1000-cycle, 20% strain stretching test [plot (a)] and the
corresponding reduction peak potentials and peak currents [plot
(b)]. Plots (c) and (d) show the CA response at -0.2 V applied
anodic potential of the PB electrode every 200 cycles in a
1000-cycle, 20% strain stretching test [plot (a)], and the
corresponding end-point currents [plot (d)]. Electrochemical tests
were performed in 0.1M PBS with pH 7.4, against an Ag/AgCl
reference electrode.
[0211] FIG. 18 shows data plots depicting sensor electrochemical
stability under repeated stretching tests. Plots (a-b) show the
response of modified lactate sensor to 2 mM lactate before (a) and
after (b) 10 cycles of 20% stretching. Plots (c-d) show the
response of modified glucose sensor to 10 mM glucose before (a) and
after (b) 10 cycles of 20% stretching. CA was recorded using a
potential of -0.2 V in 0.1 M PBS with pH 7.4.
[0212] FIG. 19 shows images depicting structural integrity of the
stretchable silver and PB/carbon ink composites. Image (a) is a
photograph of the device before and after stretched to 120%. Images
(b)-(e) depict scanning electron micrograph (SEM) images in
different magnifications of the silver trace before (b) and after
(c) 1000 cycles of 20% stretching, and of the carbon trace before
(d) and after (e) 1000 cycles of 20% stretching.
[0213] FIG. 20 shows images and data plots illustrating
electrochemical performance under mechanical deformation. Panel (a)
shows an image (i) and a data plot (ii) of the example sensor under
20% vertical strain and of the electrochemical response (e.g.,
chronoamperometry at -0.2V in PBS 0.1M, pH 7) under 20% stretching
every 200 stretching cycles until 1000 cycles (RSD=18.6%). Panel
(a) shows an image (i) and data plot (ii) of the example sensor
under 20% horizontal strain and of the electrochemical response
(e.g., chronoamperometry at -0.2V in PBS 0.1M, pH 7) under 20%
stretching every 200 stretching cycles until 1000.times.
(RSD=15.8%).
[0214] FIG. 21 shows images and data plots illustrating the
electrochemical performance of the GOx modified biosensor under
mechanical deformation. Panel (a) shows an image and data plots
depicting the example sensor under 20% vertical strain and the
electrochemical response of the sensor to 10 mM glucose after
stretching the sensor vertically 100 times at 120% and recording
the signal every 10 stretching cycles (RSD=3.45%) in data plot (i),
and the response while the sensor was under stress, after every
10-stretching deformation (RSD=5.42%) in data plot (ii). Panel (b)
shows an image and data plots depicting the example sensor under
20% horizontal strain and the electrochemical response of the
sensor to 10 mM glucose after stretching the sensor horizontally
100 times at 120% and recording the signal every 10 stretching
cycles (RSD=2.33%) in data plot (i), and the response under stress
after every 10-stretching deformation (RSD=3.14%) in data plot
(ii). Panel (c) shows a data plot depicting the variation of
current response to 10 mM glucose for all deformations (RSD=5.18%).
Panel (d) shows a data plot depicting the error associated to each
deformation cycle including all deformations; higher error is
associated to the first cycles.
[0215] FIG. 22 shows a data plot depicting the BP signal from an
example wearable acoustic-chem sensor device measured on-body while
turning the neck 90.degree. to the side, with no obvious change of
signal quality.
[0216] FIG. 23 shows data plots depicting the BP variation from an
example wearable acoustic-chem sensor device during the Valsalva
maneuver. Plot (a) depicts the event timeline of the BP signal
recording during a Valsalva maneuver. Plot(b) depicts the BP
waveform during the initial phase of the Valsalva maneuver. A
sudden increase in BP is observed. The local peaks (systolic BP)
and troughs (diastolic BP) were indicated (blue points--before
maneuver, yellow points--during maneuver). Plot (c) depicts
comparison of the systolic and diastolic BP to the initial cuff
calibration, the average BP before and during the maneuver measured
by the patch, and the validation using the cuff during the
maneuver.
[0217] FIG. 24 shows data plots depicting standard additions to
determine caffeine concentration in sweat. Plot (a) depicts
differential pulse voltammetry (DPV) of caffeine in collected sweat
(after drinking 114 mg caffeine). Increasing concentrations of
caffeine were added to the collected sweat and the respective
calibration curve was used to analyze the initial caffeine
concentration in sweat in the Plot (b) calibration curve. The
horizontal axis shows the concentration range used in the test. The
vertical axis shows the current change after each addition.
[0218] FIG. 25 shows data plots depicting the reversibility test
for the voltammetric caffeine sensor. Solutions containing only
acetate buffer 0.01M, pH 4.5 and 200 .mu.M of caffeine in the same
buffer were measured alternately. Plot (a) depicts DPV for caffeine
(green) followed by the DPV of the acetate buffer. Every time the
solution was wiped from the electrode surface, and surface rinsed
with buffer for the next measurement of buffer or caffeine. Plot
(b) depicts peak current for each caffeine measurement compared
with the buffer alone; the buffer signal was normalized to
zero.
[0219] FIG. 26 shows diagrams depicting an example embodiment of a
method 2600 for preparation and assembly of the hydrogel layers.
Diagram (a) of FIG. 26 depicts an anode hydrogel preparation
process 2610 of the method 2600. At step (i) of process 2610, the
method 2600 includes keeping a mixture of 4% agarose and DI water
under continuous stirring (e.g., at 150.degree. C.) until complete
dissolution, then adding the 2% pilocarpine nitrate under stirring.
At step (ii) of the process 2610, the method 2600 includes
depositing (e.g., drop-casting) a volume (e.g., 300 .mu.L) of the
solution in the mold (e.g., ECOFLEX molds). After cooling, for
example, the viscous solution on the mold became a solid coin shape
hydrogel. Diagram (b) of FIG. 26 depicts cathode hydrogel
preparation 2620 of the method 2600. At step (i) of the process
2620, the method 2600 includes keeping the mixture of 4% agarose
and 0.1 M PBS (pH 7.4) buffer at continuous stirring (e.g., at
150.degree. C.) until observing complete agarose dissolution. At
step (ii) of the process 2620, the method 2600 includes depositing
(e.g., drop-casting) a volume (e.g., 300 .mu.L) of the solution in
ECOFLEX molds. After cooling, for example, the viscous solution on
the mold became a solid coin shape hydrogel. Diagram (c) of FIG. 26
depicts hydrogel assembly process 2630 of the method 2600. At step
(i) of the process 2630, the method 2600 includes cutting coin
shape anode and cathode hydrogel disks with the shape of the
screen-printed pattern of the anode and cathode respectively. At
step (ii) of the process 2630, the method 2600 includes, after the
shape was provided, placing the anode hydrogel on the left side and
placing the cathode hydrogel on the right side.
[0220] FIG. 27 shows photos depicting the assembly of iontophoretic
and ultrasound hydrogels for an example embodiment of a wearable
acoustic-electrochemical sensor device. Photo (a) depicts an
example commercial solid gel pad for ultrasound inspection
integrated on device. The picture shows a freestanding cut piece of
solid gel. The hydrogels for cathode and anode were cut into the
shape of the IP electrodes. Photo (b) depicts the example solid gel
pad after the shape was provided, and photo (c) depicts the anode
hydrogel was placed on the left side and the cathode hydrogel was
placed in the right side.
[0221] FIG. 28 shows photos depicting a transfer process of the
example wearable acoustic-electrochemical sensor device of FIG. 27.
Photo (a) depicts double-sided tattoo adhesive. Photo (b) depicts
opening for the sensing areas. Photo (c) depicts removing the first
protective layer from the double-sided tattoo adhesive with opening
for the sensing areas. Photo (d) depicts applying adhesive to the
tattoo; Photo (e) depicts after removing the second protective
layer from the applied adhesive. Photo (f) depicts placing the
hydrogels and US gel. Photos (g) and (h) depict transferring to the
body.
[0222] FIG. 29 shows a diagram and data plots depicting an example
characterization of an example multimodal wearable sensor. Diagram
(a) depicts signal interference study between the sweat caffeine
electrochemical sensor (left side) and the BP transducer, including
data plot (b) depicting BP signal recording while initially
sweeping the potential for caffeine detection followed by
terminating the sweeping (off), and data plot (c) depicting BP
signal recording while the potential sweeping for caffeine
detection is off followed by initiating the sweeping (on). The
effect of the BP signal on the caffeine detection was also
investigated. Data plot (d) depicts DPV was recording for caffeine
detection while the BP was active, following by terminating the BP
signal acquisition (off). Data plot (e) depicts DPV recording for
caffeine detection while the BP was inactive, following by
initiating the BP signal acquisition (on).
[0223] FIG. 30 shows diagrams and data plots illustrating an
example on body cross-talking evaluation of an example multimodal
wearable sensor. Diagram (a) depicts signal interference study
between iontophoretic current (0.3 mA/cm2) applied between cathode
and anode (dark orange color) and blood pressure device (light
orange color). Data plot (b) depicts blood pressure wave form
recorded when the IP current was initially on, and next turned off.
Data plot (c) depicts blood pressure waveform recorded when the IP
current was initially off and then turned on. Diagram (d) depicts
cross talking study between detection potential for glucose (dark
green) and the blood pressure ultrasound device (light green). Data
plot (e) depicts BP signal recorded while the detection potential
was initially off, followed by turning the detection potential
(-0.2V) on (red dotted line) and off (blue dotted line). Data plot
(f) and (g) depict glucose signal recording while applying an
on/off ultrasound cycle every 60 seconds during 3 min. Diagram (h)
depicts cross talking study between the lactate detection potential
(dark blue) and the blood pressure ultrasound device (light blue).
Data plot (i) depicts BP signal recorded while the detection
potential was initially off, followed by turning the detection
potential -0.2V on (red dotted line) and off (blue dotted line).
Data plot (j) depicts lactate signal recording while applying the
ultrasound cycle on/off every 60 seconds during 3 min. Data plot
(k) depicts lactate signal recording while applying the ultrasound
cycle off/on every 60 seconds during 3 min.
[0224] FIG. 31 shows diagrams and data plots depicting an example
in vitro cross-talking evaluation of an example multimodal wearable
sensor. In vitro measurements were performed using the agarose-PBS
hydrogel and the ultrasound gel placed over the blood pressure
transducers. Panel (a) shows a diagram and data plot depicting the
PBS hydrogel and ultrasound gel were in contact while the
ultrasound pulse was turned on and off (right), respectively. A
decreasing (more positive current) was observed when the gels were
in contact. Panel (b) shows a diagram and data plot depicting a
different design with increased distance between the gels were used
in vitro with the same amount of ultrasound gel, due to the
physical distance, no cross talking was observed (right). Panel (c)
shows a diagram and data plot depicting the same design (with
shorter distance between the blood pressure and chemical sensors)
was used with a solid hydrogel and no apparent cross talking was
observed (right).
[0225] FIG. 32 shows images and data plots depicting an example
implementation for preparing and testing an example solid
ultrasound gel. Image (a) depicts the commercial solid gel pad for
ultrasound inspection. Image (b) depicts the razor blades used to
cut thin slice of solid ultrasound gel; the insert shows the gap
between the blades are 800 Image (c) depicts integrated solid gel
on device. The insert in the image (c) shows a freestanding cut
piece of solid gel. Panel (d) includes data plots depicting
ultrasound penetration intensity test with a pulse-echo test that
is performed in data plot (i) using a phantom ECOFLEX with the
liquid and solid ultrasound gel and in the absence of the gel; the
respective echo amplitude is compared in data plot (ii). Panel (e)
includes data plots depicting on body experiment comparing the BP
waveform measured with liquid [data plot (i)] and solid ultrasound
gel [data plot (ii)].
[0226] FIG. 33 shows a diagram and data plots depicting the effect
of an example substrate for a wearable, integrated
acoustic-electrochemical sensor device on ultrasound transmission.
Substrates with certain ultrasound impedance will result in a
different ultrasound penetration intensity. Diagram (a) depicts an
illustration of an example fabrication process. A pulse-echo test
was performed on different materials as substrates, including
Thermoplastic Polyurethane (TPU), polyurethane (PU),
styrene-ethylene-butylene-styrene block copolymer (SEBS), and
ECOFLEX. Data plot (b) shows example data of pulse echoes for TPU
(plot (i)) and ECOFLEX (plot (ii)); data plot (c) shows example
data of pulse echoes for PU (plot (i)) and ECOFLEX (plot (ii)); and
data plot (d) shows example data of pulse echoes for SEBS (plot
(i)) and ECOFLEX (plot (ii)). The echo signal intensity is directly
compared for the ultrasound penetration.
[0227] FIG. 34 shows data plots depicting on body evaluation for
caffeine intake. Data plots in panel (a) depict an on-body
evaluation of BP changes for a volunteer with no habitual caffeine
intake (caffeine intolerant), before and after caffeine sugar free
beverage consumption (right), and the bar graph represent the
sensor validation using a commercial cuff (white) and BP readings
obtained with ultrasound transducer (orange). Data plots in panel
(b) depict an on-body evaluation of BP changes for a volunteer with
regular caffeine intake habits (caffeine tolerant), before and
after caffeine sugar free beverage consumption (right), and the bar
graph represent the sensor validation using a commercial cuff
(white) and BP readings obtained with ultrasound transducer
(orange).
[0228] FIG. 35 shows diagrams and data plots illustrating on-body
caffeine detection and pH variation. Panel (a) shows a diagram and
data plots (i) and (ii) that show a gel configuration for caffeine
detection in stimulated sweat without the use of acetate buffer
over the caffeine sensor; the baseline before caffeine intake
presents no peak (data plot(i)), while the caffeine peak is
presented at 1.6V when the caffeine sensor is in direct contact
with sweat (.about.pH 7) (data plot (ii)). Panel (b) shows a
diagram and data plots (i) and (ii) that show the gel configuration
using an acetate buffer pH 4.5 loaded agarose gel on the caffeine
sensor (green gel); the baseline before caffeine intake presents no
peak (data plot(i)), while a defined caffeine peak is presented at
1.2V when the caffeine sensor is in contact with the acetate buffer
gel (data plot(ii)).
[0229] FIG. 36 shows data plots illustrating control experiments:
electrochemical sensors without the sensing recognition layer. Plot
(a) depicts the cathode sensor signal recorded after the first ISF
extraction (red line) and after the second ISF extraction (black
line). Plot (b) depicts the anode sensor signal recorded in the
first sweat stimulation (red line) and during the second sweat
stimulation (black line). The applied potential was -0.2 V.
[0230] FIG. 37 shows data plots depicting control experiments:
response for lactate and glucose recording without exercise and
food ingestion. Plot (a) depicts the lactate sensor response after
the first sweat stimulation (green dash line) and second sweat
stimulation (green solid line). Plot (b) depicts the glucose sensor
response after the first ISF extraction (red dash line) and second
ISF extraction (red solid line). The applied potential was -0.2
V.
[0231] FIG. 38 shows data plots depicting control experiments:
response for alcohol and glucose recording without alcohol and food
ingestion. Plot (a) depicts the alcohol the first sweat stimulation
(blue dash line) and second sweat stimulation (blue solid line).
Plot (b) depicts the response of the glucose sensor after the first
ISF extraction (red dash line) and second ISF extraction (red solid
line). The applied potential was -0.2 V.
[0232] FIG. 39A shows an image depicting the design of an epidermal
sensor patch for the simultaneous monitoring of blood pressure
along with sweat alcohol, caffeine and lactate, and ISF glucose
chemical markers. The example epidermal sensor patch of FIG. 39A
can be configured similar to the example embodiment of the wearable
acoustic-chem sensor device 100J, shown in FIG. 1J. For example,
the epidermal sensor patch can include a blood pressure sensor
comprising an array of ultrasonic transducers, and a plurality of
electrochemical sensors to non-invasively and electrochemically
detect biomarker levels from biofluids, such as sweat and/or
interstitial fluid (ISF). The example physiological sensor and
electrochemical sensors of the epidermal sensor patch are in
electrical connection with electrical contact sites via electrical
interconnections. The example epidermal sensor patch is configured
for simultaneous monitoring of four analytes: glucose ("G"),
lactate ("L"), caffeine ("C"), and alcohol ("A") at four distinct
working electrodes of the electrochemical sensors.
[0233] In some implementations, for example, the epidermal sensor
patch of FIG. 39A can be used for detection and/or monitoring of
sepsis in a patient user. Sepsis is a leading cause of acute
hospital mortality, e.g., affecting more than 30 million people
worldwide every year and causing approximately six million deaths.
In the United States, sepsis is estimated to cost nearly $24
billion, which is the most expensive condition presently treated in
the U.S. hospital environment. Sepsis can involve two components:
unrecovered arterial hypotension (e.g., with mean arterial pressure
(MAP) of less than 70 mmHg) and hyperlactatemia (e.g., with lactate
level greater than 2 mmol/L). These two components can persist
despite adequate fluid resuscitation. Yet, sepsis is potentially
preventable and treatable if identified early with rapid treatment
initiation. Studies have shown that early sepsis detection can
impact overall survival, as survival decreases with every hour
delay in initiation of effective antimicrobials.
[0234] Existing techniques and technologies for monitoring sepsis
are limited and difficult to implement, let alone lacking a
singular, integrated device that monitors multiple biomarkers of
sepsis in a non-intrusive, non-invasive manner. For instance,
sepsis monitoring can include common blood pressure monitoring
using the peripheral cuff that straps around a user's arm and
inflates for a singular measurement; since this technique is slow
and uncomfortable to users, it is typically limited to one blood
pressure measurement every five minutes. Also, for instance, to
obtain measurements for analyte markers, such as lactate, sodium,
glucose or other, current techniques involve the highly invasive
blood draw and/or an arterial catheter.
[0235] FIG. 39B shows another diagram of the example epidermal
sensor patch for continuous and simultaneous monitoring of sepsis
by blood pressure sensing coupled with electrochemical sensing of
three analytes: lactate in sweat, sodium in sweat, and glucose in
ISF.
[0236] FIG. 39C shows a diagram illustrating the sensing mechanism,
e.g., for monitoring of sepsis, using the example epidermal sensor
patch of FIG. 39B. For example, the electrochemical sensor for
monitoring of lactate from sweat can implement amperometry based on
a LOx-modified working electrode (e.g., see FIG. 10, panel (a))
proximate a counter electrode and reference electrode; and
iontophoretic electrode can be used to assist in stimulating the
sweat from the skin. Also, for example, the electrochemical sensor
for monitoring of sodium (or other electrolytes) from sweat can
implement potentiometry using an example ISE working electrode and
Ag/AgCl reference electrode. Further, for example, the
electrochemical sensor for monitoring of glucose from ISF can
implement an iontophoretic current to extract ISF from skin and
amperometry for detection of glucose in the ISF based on a
GOx-modified working electrode (e.g., see FIG. 10, panel (b))
proximate a counter electrode and reference electrode. Moreover,
for example, the blood pressure sensor can include an array of
ultrasound transducers that generate ultrasound signals transmitted
toward a pulsating artery, where movement of the anterior wall
(ant-wall) and posterior wall (post-wall) of the artery can cause
return ultrasound signals (echoes), where peaks shifts are used to
determine one or more parameters associated with blood
pressure.
[0237] FIG. 39D shows data plots depicting example amperometric
current data of lactate sensing obtained from an example
implementation of sepsis monitoring of a subject using the example
epidermal sensor patch of FIG. 39B.
[0238] FIG. 39E shows data plots depicting example ultrasound
signal data obtained from an example implementation of sepsis
monitoring of a subject using the example epidermal sensor patch of
FIG. 39B.
[0239] FIG. 39F shows data plots depicting example amperometric
current data of glucose sensing obtained from an example
implementation of sepsis monitoring of a subject using the example
epidermal sensor patch of FIG. 39B.
[0240] FIG. 39G shows data plots depicting example data of an
on-body implementation of an example epidermal sensor patch, as in
FIG. 39B, which monitored the subject's blood pressure and at least
one analyte (e.g., lactate) before, during and after exercise.
[0241] FIG. 40 shows an image and an illustrative diagram depicting
the design of an example embodiment of a stretchable integrated
blood pressure-electrochemical sensing patch device for the
simultaneous detection of sweat sodium and lactate, and blood
pressure. Image (a) is a photo image of the sensor on the body.
Diagram (b) is an illustration depicting the example sensor's
acoustic and electrochemical sensing components. Lactate and sodium
sensors are located at the cathodic compartment. A three-electrodes
system is used for lactate detection (red circle) and a
two-electrode system is used for sodium detection (blue square).
The blood pressure sensor is located in the center on the patch
(black square).
[0242] FIG. 41 shows data plots depicting in vitro characterization
of a sodium sensor in example implementations. Plot (a) depicts the
reversibility test for the sequential increasing and decreasing
NaCl concentrations in a single sensor. The reversibility was
realized for two sensors (red and black curves) using i, 0.1 mM,
ii, 1 mM, iii, 10 mM and iv, 100 mM of NaCl. Plot (b) shows a
calibration curve for the response of the sodium sensor to
concentrations i-iv. (n=5, RSD=5%, Slope=0.73, r2=0.99).
[0243] FIG. 42 shows data plots depicting another in vitro
characterization of a sodium sensor in example implementations.
Plot (a) depicts the interference test comparing the potentiometric
sensor response to NaCl vs KCl, using increasing concentrations of
0.1, 1, 10, and 100 mM of NaCl (black curve) and KCL (red curve).
Plot (b) depicts the stability of the sodium potentiometric
response during a continuous 75 min monitoring of 0.1 mM NaCl.
[0244] FIG. 43 shows a data plot depicting an in vitro
characterization of a lactate sensor in the presence of sodium in
example implementations. Effect of sodium upon the amperometric
response of the lactate sensor. A stable PBS baseline was acquired
at -0.2V (dotted line) followed by the addition of 5 mM lactate
(blue curve), and subsequent addition of 20 mM NaCl (red
curve).
[0245] FIG. 44 shows data plots depicting continuous
sodium/lactate/BP/HR performance. Data plots (a) depict continuous
signal recording showing sweat sodium profile [data plot (i)] and
lactate profile [data plot (ii)] during stationary biking for fit
subject, and depict BP/HR signal recording before (green), during
(red), and after (purple) stationary biking [data plot (iii)]. Data
plots (b) depict continuous signal recording showing sweat sodium
profile [data plot (i)] and lactate profile [data plot (ii)] during
stationary biking for a sedentary subject, and depict BP/HR signal
recording before (green), during (red), and after (purple)
stationary biking [data plot (iii)].
[0246] FIG. 45 shows data plots depicting continuous alcohol
monitoring in stimulated sweat for two volunteers. Plots (a) and
(b) depict continuous alcohol monitoring was performed by measuring
alcohol levels in sweat every 10 minutes. Sweat was stimulated
before drinking alcohol by performing 10 minutes IP, followed by 5
minutes waiting time for sweat generation. Chronoamperometry was
performed and the fifth amperogram was taken (i). Therefore, the
total time for the final signal was 30 minutes. After every sweat
stimulation a breathalyzer was used to measure BAC. Sweat was
stimulated every 10 minutes until .about.zero BAC, and the
correlation between sweat alcohol (black plot) and blood alcohol
(blue plot) is shown in the bottom plot (ii).
[0247] FIG. 46 shows data plots depicting correlation curves for
sweat and ISF analytes. Plot (a) depicts the correlation curve for
ISF glucose and blood glucose (n=13). Plot (b) depicts the
correlation curve for sweat lactate and blood lactate=32 (n=18).
Plot (c) depicts a correlation curve for sweat alcohol and blood
alcohol (n=10).
Examples
[0248] In some embodiments in accordance with the present
technology (example A1), a sensor device includes one or more
ultrasound sensors for sensing a physiological characteristic; and
one or more electrochemical sensors for sensing a chemical
characteristic, wherein the ultrasound sensor and the
electrochemical sensor share a substrate.
[0249] Example A2 includes the sensor device as in any of examples
of A1-A14, wherein the one or more ultrasound sensors measure one
or more of a blood pressure or a heart rate.
[0250] Example A3 includes the sensor device as in any of examples
of A1-A14, further comprising a temperature sensor.
[0251] Example A4 includes the sensor device as in any of examples
of A1-A14, further comprising an electrocardiogram (ECG)
sensor.
[0252] Example A5 includes the sensor device as in any of examples
of A1-A14, further comprising a pressure sensor.
[0253] Example A6 includes the sensor device as in any of examples
of A1-A14, further comprising a mechanical strain sensor.
[0254] Example A7 includes the sensor device as in any of examples
of A1-A14, further comprising one or more additional sensors
including a temperature sensor; an electrocardiogram (ECG) sensor;
a pressure sensor; and/or a mechanical strain sensor.
[0255] Example A8 includes the sensor device as in any of examples
of A1-A14, wherein the one or more electrochemical sensors measure
one or more of a sodium concentration; a potassium concentration; a
glucose concentration; and/or a lactate concentration.
[0256] Example A9 includes the sensor device as in any of examples
of A1-A14, wherein the ultrasound sensor includes a blood pressure
sensor and the electrochemical sensor includes a lactate sensor,
wherein the sensor monitors for septic shock.
[0257] Example A10 includes the sensor device as in any of examples
of A1-A14, wherein the ultrasound sensor includes a blood pressure
sensor.
[0258] Example A11 includes the sensor device as in any of examples
of A1-A14, wherein the electrochemical sensor includes a lactate
sensor.
[0259] Example A12 includes the sensor device as in any of examples
of A1-A14, wherein the substrate includes one or more of:
ECOFLEX.RTM.; polydimethylsiloxane (PDMS); and/or polyurethane.
[0260] Example A13 includes the sensor device as in any of examples
of A1-A14, wherein the sensor device is structured to be a wearable
or attachable to a user.
[0261] Example A14 includes the sensor device as in example A12 or
any of examples of A1-A13, wherein the wearable sensor is worn on
the epidermis.
[0262] In some embodiments in accordance with the present
technology (example B1), a wearable, acoustic-electrochemical
sensor device includes a flexible substrate comprising an
electrically insulative material, the flexible substrate capable of
attaching and conforming to skin; an electrochemical sensor
comprising two or more electrodes disposed on the flexible
substrate, the two or more electrodes including a first electrode
to detect a signal associated with an analyte by an electrochemical
detection, and a second electrode configured as a counter electrode
or a reference electrode; a physiological sensor comprising an
array of acoustic transducers disposed on the flexible substrate
and a ground wire coupled to and spanning across each acoustic
transducer of the array, the array of acoustic transducers
including an acoustic transduction material, wherein the
physiological sensor is configured to direct acoustic signals from
the array of acoustic transducers toward a blood vessel in or
beneath the skin to detect a hemodynamic parameter of the blood
vessel; and an array of electrical interconnection structures
disposed on the flexible substrate, wherein at least one of the
electrical interconnection structures is configured as a ground
electrical interconnection structure, and wherein the ground wire
of the physiological sensor spans from the array of acoustic
transducers to the ground electrical interconnection structure,
wherein the sensor device is operable to simultaneously detect and
monitor one or more analyte markers and physiological markers.
[0263] Example B2 includes the sensor device of any of examples
B1-B20, wherein the array of acoustic transducers of the
physiological sensor is spaced apart from the electrochemical
sensor by a distance of at least 0.1 cm.
[0264] Example B3 includes the sensor device of any of examples
B1-B20, wherein the physiological sensor is configured on a first
side of the flexible substrate configured to attach to the skin,
and the electrochemical sensor is configured on a second side of
the flexible sensor opposite to the first side, such that the
electrochemical sensor is able to be exposed to a biofluid
deposited on the electrochemical sensor.
[0265] Example B4 includes the sensor device of any of examples
B1-B20, wherein the physiological sensor includes a hydrogel
material coupled to the array of acoustic transducers and
configured to propagate an acoustic signal generated at the
acoustic transducers to the skin and to propagate a returned
acoustic echo received from the skin to the acoustic
transducers.
[0266] Example B5 includes the sensor device of any of examples
B1-B20, wherein the electrochemical sensor includes a
functionalization layer disposed at least partially on the first
electrode that includes one or more molecules to catalyze a
chemical reaction or bind to the analyte for the electrochemical
detection at the first electrode, and wherein the wearable,
acoustic-electrochemical sensor device further comprises: a second
electrochemical sensor comprising two or more electrodes disposed
on the flexible substrate, the two or more electrodes of the second
electrochemical sensor including a third electrode to detect a
second signal associated with a second analyte by a second
electrochemical detection, and a fourth electrode configured as a
counter electrode or a reference electrode, wherein the second
analyte is different than the analyte detectable at the first
electrode.
[0267] Example B6 includes the sensor device of any of examples
B1-B20, wherein the second electrode is configured as the reference
electrode, and wherein the two or more electrodes of the
electrochemical sensor include a third electrode configured as the
counter electrode.
[0268] Example B7 includes the sensor device of example B6 or any
of examples B1-B20, wherein the two or more electrodes of the
electrochemical sensor include a fourth electrode configured as an
iontophoresis (IP) electrode, the IP electrode operable to
facilitate extraction of interstitial fluid of the skin or induce
excretion of sweat from the skin.
[0269] Example B8 includes the sensor device of example B7 or any
of examples B1-B20, wherein the electrochemical sensor includes a
hydrogel coupled to the IP electrode, wherein the hydrogel entraps
one or more chemicals able to cause extraction of the interstitial
fluid or excretion of the sweat upon controlled release from the
hydrogel by an electrical potential applied at the IP
electrode.
[0270] Example B9 includes the sensor device of example B6 or any
of examples B1-B20, wherein two or more electrodes are printed
electrodes, wherein the first electrode and the counter electrode
comprise a Prussian Blue, and wherein the reference electrode
comprise a silver ink.
[0271] Example B10 includes the sensor device of any of examples
B1-B20, wherein the electrical interconnection structures are
configured as serpentine interconnection structures that allow for
stretching and bending on the flexible substrate.
[0272] Example B11 includes the sensor device of any of examples
B1-B20, wherein the acoustic transduction material includes at
least one of piezoelectric lead zirconate titanate (PZT), lead
magnesium niobate-lead titanate (PMN-PT), or polyvinylidene
difluoride (PVDF).
[0273] Example B12 includes the sensor device of example B11 or any
of examples B1-B20, wherein each transducer pixel includes an
aspect ratio of 0.3 or smaller based on a height dimension to a
width dimension, such that aspect ratio is able to control
vibration of the acoustic transduction material to be in a
thickness mode with a particular frequency or frequency range.
[0274] Example B13 includes the sensor device of example B12 or any
of examples B1-B20, wherein the particular frequency is 7 MHz; or
wherein the frequency range includes 5 MHz to 9 MHz.
[0275] Example B14 includes the sensor device of any of examples
B1-B20, wherein the flexible substrate includes at least one of a
styrene-ethylene-butylene-styrene block copolymer (SEBS), a
styrene-isoprene-styrene block copolymer (SIS), or a
styrene-butylene-styrene block copolymer (SBS).
[0276] Example B15 includes the sensor device of any of examples
B1-B20, wherein the flexible substrate includes at least one of
ECOFLEX.RTM., polydimethylsiloxane (PDMS), thermoplastic
polyurethane (TPU), polyurethane (PU), or polyethylene vinyl
acetate (PEVA).
[0277] Example B16 includes the sensor device of any of examples
B1-B20, wherein the flexible substrate is structured to include a
first substrate layer and a second substrate layer that is attached
to a side of the first substrate layer, wherein each of the first
substrate layer and the second substrate layer comprises a first
region and a second region, wherein the physiological sensor is
coupled to the first region of the first substrate layer, and the
electrochemical sensor is coupled to the second region of the
second substrate layer, wherein the second substrate layer includes
an opening at the first region such that physiological sensor is
exposed through the opening of the second substrate layer.
[0278] Example B17 includes the sensor device of any of examples
B1-B20, wherein the hemodynamic parameter includes blood pressure
or blood flow.
[0279] Example B18 includes the sensor device of any of examples
B1-B20, further comprising one or more additional sensors including
a temperature sensor, an electrocardiogram (ECG) sensor, a pressure
sensor, or a mechanical strain sensor.
[0280] Example B19 includes the sensor device of any of examples
B1-B20, wherein the physiological sensor comprising the acoustic
transducers is operable to detect blood pressure of a user of the
wearable, acoustic-electrochemical sensor device, and wherein the
electrochemical sensor is operable to detect lactate of the user,
such that the sensor device is operable to monitor for septic
shock.
[0281] Example B20 includes the sensor device of any of examples
B1-B19, wherein the analyte includes lactate, cortisol, glucose,
alcohol, caffeine, or an electrolyte.
[0282] In some embodiments in accordance with the present
technology (example C1), a method for fabricating a wearable,
acoustic-electrochemical sensor device includes providing a
flexible substrate; producing a pattern of electrode structures on
the flexible substrate; producing a pattern of electrical
interconnection structures on the flexible substrate; producing a
pattern of electrical contact site structures on the flexible
substrate, wherein the electrical interconnection structures are
each coupled between an electrode structure of the pattern of
electrode structures and an electrical contact site structure of
the pattern of electrical contact site structures; producing an
electrically insulating layer over the flexible substrate covering
the pattern of electrical interconnection structures and without
covering at least a portion of each electrode structure and at
least a portion of each electrical contact site structure; and
producing an acoustic transducer on the flexible substrate by:
flipping the flexible substrate backside up and producing a second
pattern of electrical interconnection structures and a second
pattern of electrical contact site structures on the backside of
the flexible substrate, soldering (e.g., adhering, bonding) a
pattern of transducer chip structures, by using a conductive ink
solvent as an adhesive to bond a structure, at terminuses of the
electrical interconnection structures, wherein at least one of the
electrical interconnection structures of the second pattern does
not include a soldered transducer chip structure so as to serve as
a reserved ground interconnect, and soldering (e.g., adhering,
bonding) an electrical ground structure, using a solvent, to a
first side of the transducer chip structures that spans across the
pattern of transducer chip structures to the reserved ground
interconnect.
[0283] Example C2 includes the method of any of examples C1-C14,
wherein the flexible substrate includes at least one of a
styrene-ethylene-butylene-styrene block copolymer (SEBS), a
styrene-isoprene-styrene block copolymer (SIS), a
styrene-butylene-styrene block copolymer (SBS),
polydimethylsiloxane (PDMS), thermoplastic polyurethane (TPU),
polyurethane (PU), polyethylene vinyl acetate (PEVA), or
ECOFLEX.RTM..
[0284] Example C3 includes the method of any of examples C1-C14,
wherein the producing the pattern of electrode structures includes
printing the electrode structures using an electrically conductive
ink material.
[0285] Example C4 includes the method of example C3 or any of
examples C1-C14, wherein the printing the electrode structures
includes using one or both of a silver ink and Prussian Blue ink to
print the pattern of electrode structures over the flexible
substrate.
[0286] Example C5 includes the method of any of examples C1-C14,
wherein the electrode structures include at least one working
electrode and one or more of (i) at least one of an iontophoresis
(IP) electrode, (ii) at least one of a reference electrode, or
(iii) at least one of a counter electrode.
[0287] Example C6 includes the method of example C5 or any of
examples C1-C14, wherein the at least one of the IP electrode
and/or the at least one of the reference electrode is printed using
a silver ink.
[0288] Example C7 includes the method of example C5 or any of
examples C1-C14, wherein the at least one working electrode and/or
the at least one of the counter electrode is printed using a
Prussian Blue ink.
[0289] Example C8 includes the method of any of examples C1-C14,
wherein the producing the pattern of electrical interconnection
structures includes printing serpentine interconnection structures
using an electrically conductive ink, wherein the printed
serpentine interconnection structures allow for stretching and
bending of the pattern of electrical interconnection structures on
the flexible substrate.
[0290] Example C9 includes the method of example C8 or any of
examples C1-C14, wherein the serpentine interconnection structures
are printed using a silver ink.
[0291] Example C10 includes the method of any of examples C1-C14,
wherein the producing an electrically insulating layer includes
printing a resin comprising a styrene-ethylene-butylene-styrene
block copolymer (SEBS) material.
[0292] Example C11 includes the method of any of examples C1-C14,
further comprising: chemically modifying an exposed surface of at
least one electrode structure to configure an electrochemical
sensor for sensing a target analyte.
[0293] Example C12 includes the method of any of examples C1-C14,
wherein the producing the second pattern of electrical
interconnection structures includes printing serpentine
interconnection structures using an electrically conductive ink,
wherein the printed serpentine interconnection structures allow for
stretching and bending of the second pattern of electrical
interconnection structures on the flexible substrate.
[0294] Example C13 includes the method of example C12 any of
examples C1-C14, wherein the serpentine interconnection structures
are printed using a silver ink.
[0295] Example C14 includes the method of any of examples C1-C13,
wherein the soldering the transducer chips includes depositing an
organic solvent by dripping droplets of the organic solvent on an
interface between the terminuses of the electrical interconnection
structures and the transducer chips.
[0296] In some embodiments in accordance with the present
technology (example D11), a method for fabricating an acoustic
transducer for a wearable acoustic sensor device includes:
providing a flexible substrate; producing a pattern of electrical
interconnection structures and a pattern of electrical contact site
structures on the flexible substrate, wherein the electrical
interconnection structures are coupled to the electrical contact
site structures at one end; depositing an organic solvent at a
terminus location of the electrical interconnection structures to
partially dissolve the electrically conductive material of the
electrical interconnection structures at the terminus location,
wherein at least one of the electrical interconnection structures
does not receive the deposited organic solvent, such that the at
least one of the electrical interconnection structures that does
not receive the deposited organic solvent is to serve as a ground
interconnect; producing an array of acoustic transducer structures
coupled to the pattern of electrical interconnection structures,
where, for each acoustic transducer structure, placing an acoustic
transducer structure at the terminus location to allow bonding of
the acoustic transducer structure to a respective electrical
interconnection structure; depositing an organic solvent at a
portion of a surface of the acoustic transducer structures to
partially dissolve an acoustic transduction material of the
acoustic transducer structures; and attaching a wire across each of
the acoustic transducer structures to connect to the ground
interconnect.
[0297] Example D2 includes the method of any of examples D1-D6,
wherein the acoustic transduction material includes at least one of
piezoelectric lead zirconate titanate (PZT), lead magnesium
niobate-lead titanate (PMN-PT), or polyvinylidene difluoride
(PVDF).
[0298] Example D3 includes the method of any of examples D1-D6,
wherein the producing an array of acoustic transducer structures
includes depositing a bonding material for bonding with the
acoustic transducer structure to be placed.
[0299] Example D4 includes the method of example D3 or any of
examples D1-D6, wherein the bonding material includes an
electrically conductive ink.
[0300] Example D5 includes the method of any of examples D1-D6,
wherein the depositing the organic solvent includes dripping
droplets of the organic solvent on the terminus location.
[0301] Example D6 includes the method of any of examples D1-D6,
wherein the organic solvent includes one or more of toluene,
xylene, benzene, cyclopentane, n-pentane, cyclohexane,
cyclohexanone, ethylbenzene, acetone, methanol, ethanol,
isopropanol, tetrahydrofuran, dimethyl sulfoxide, or the mixture
thereof.
[0302] Implementations of the subject matter and the functional
operations described in this patent document can be implemented in
various systems, digital electronic circuitry, or in computer
software, firmware, or hardware, including the structures disclosed
in this specification and their structural equivalents, or in
combinations of one or more of them. Implementations of the subject
matter described in this specification can be implemented as one or
more computer program products, i.e., one or more modules of
computer program instructions encoded on a tangible and
non-transitory computer readable medium for execution by, or to
control the operation of, data processing apparatus. The computer
readable medium can be a machine-readable storage device, a
machine-readable storage substrate, a memory device, a composition
of matter effecting a machine-readable propagated signal, or a
combination of one or more of them. The term "data processing unit"
or "data processing apparatus" encompasses all apparatus, devices,
and machines for processing data, including by way of example a
programmable processor, a computer, or multiple processors or
computers. The apparatus can include, in addition to hardware, code
that creates an execution environment for the computer program in
question, e.g., code that constitutes processor firmware, a
protocol stack, a database management system, an operating system,
or a combination of one or more of them.
[0303] A computer program (also known as a program, software,
software application, script, or code) can be written in any form
of programming language, including compiled or interpreted
languages, and it can be deployed in any form, including as a
stand-alone program or as a module, component, subroutine, or other
unit suitable for use in a computing environment. A computer
program does not necessarily correspond to a file in a file system.
A program can be stored in a portion of a file that holds other
programs or data (e.g., one or more scripts stored in a markup
language document), in a single file dedicated to the program in
question, or in multiple coordinated files (e.g., files that store
one or more modules, sub programs, or portions of code). A computer
program can be deployed to be executed on one computer or on
multiple computers that are located at one site or distributed
across multiple sites and interconnected by a communication
network.
[0304] The processes and logic flows described in this
specification can be performed by one or more programmable
processors executing one or more computer programs to perform
functions by operating on input data and generating output. The
processes and logic flows can also be performed by, and apparatus
can also be implemented as, special purpose logic circuitry, e.g.,
an FPGA (field programmable gate array) or an ASIC (application
specific integrated circuit).
[0305] Processors suitable for the execution of a computer program
include, by way of example, both general and special purpose
microprocessors, and any one or more processors of any kind of
digital computer. Generally, a processor will receive instructions
and data from a read only memory or a random access memory or both.
The essential elements of a computer are a processor for performing
instructions and one or more memory devices for storing
instructions and data. Generally, a computer will also include, or
be operatively coupled to receive data from or transfer data to, or
both, one or more mass storage devices for storing data, e.g.,
magnetic, magneto optical disks, or optical disks. However, a
computer need not have such devices. Computer readable media
suitable for storing computer program instructions and data include
all forms of nonvolatile memory, media and memory devices,
including by way of example semiconductor memory devices, e.g.,
EPROM, EEPROM, and flash memory devices. The processor and the
memory can be supplemented by, or incorporated in, special purpose
logic circuitry.
[0306] It is intended that the specification, together with the
drawings, be considered exemplary only, where exemplary means an
example. As used herein, the singular forms "a", "an" and "the" are
intended to include the plural forms as well, unless the context
clearly indicates otherwise. Additionally, the use of "or" is
intended to include "and/or", unless the context clearly indicates
otherwise.
[0307] While this patent document contains many specifics, these
should not be construed as limitations on the scope of any
invention or of what may be claimed, but rather as descriptions of
features that may be specific to particular embodiments of
particular inventions. Certain features that are described in this
patent document in the context of separate embodiments can also be
implemented in combination in a single embodiment. Conversely,
various features that are described in the context of a single
embodiment can also be implemented in multiple embodiments
separately or in any suitable subcombination. Moreover, although
features may be described above as acting in certain combinations
and even initially claimed as such, one or more features from a
claimed combination can in some cases be excised from the
combination, and the claimed combination may be directed to a
subcombination or variation of a subcombination.
[0308] Similarly, while operations are depicted in the drawings in
a particular order, this should not be understood as requiring that
such operations be performed in the particular order shown or in
sequential order, or that all illustrated operations be performed,
to achieve desirable results. Moreover, the separation of various
system components in the embodiments described in this patent
document should not be understood as requiring such separation in
all embodiments.
[0309] Only a few implementations and examples are described and
other implementations, enhancements and variations can be made
based on what is described and illustrated in this patent
document.
* * * * *